Silk fibroin materials and use thereof

ABSTRACT

The present application provides a composition comprising porous silk fibroin scaffold material. The porous silk fibroin scaffold can be used for tissue engineering. The porosity of the silk fibroin scaffold described herein can be adjusted to mimic the gradient of densities found in natural tissue.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation Application of U.S. non-provisionalpatent application Ser. No. 10/541,182 filed 14 Jun. 2006, now U.S. Pat.No. 7,842,780 B2, which is a U.S. §371 National entry of InternationalApplication number PCT/US04/00255 filed Jan. 7, 2004, which claims thebenefit of priority under 35 U.S.C. §119 (e) of U.S. ProvisionalApplication No. 60/438,393 filed Jan. 7, 2003.

GOVERNMENT SUPPORT

This invention was made with United States Government support underGrant No. 1R01 DE13405-01A1 awarded by the National Institutes ofHealth, and Grant No. DMR-0090384 awarded by the National ScienceFoundation. The United States Government has certain rights in theinvention.

SEQUENCE LISTING

The instant application contains a Sequence Listing which has beensubmitted in ASCII format via EFS-Web and is hereby incorporated byreference in its entirety. Said ASCII copy, created on Jul. 12, 2010, isnamed 20100715.txt and is 2,254 bytes in size.

FIELD OF THE INVENTION

The present invention provides methods for the production of3-dimensional porous silk fibroin scaffolds that can be used in tissueengineering. The silk fibroin scaffolds described herein areparticularly suited for tissue engineering as the porosity of thescaffold can be adjusted throughout mimicking the gradient of densitiesfound in natural tissue. Methods for producing 3-dimensional tissueusing the silk based scaffolds are also provided.

BACKGROUND OF THE INVENTION

A major goal of tissue engineering is to develop a biologicalalternative in vitro for regenerative tissue growth in vivo within adefect area. Porous polymer scaffolds play a crucial role in thethree-dimensional growth and formation of new tissue in the field oftissue engineering.

In recent years biodegradable polymers such as poly (glycolic acid),poly (L-lactic acid) (PLLA) and their copolymerspoly(L-lactic-co-glycolic acid) (PLGA) have been used as scaffoldmaterials in studies of tissue formation. (Sofia et al. Journal ofBiomedical materials research 2001, 54, 139-148). The advantages ofthese polymers is their biocompatibility and degradability. However,PLGA can induce inflammation due to the acid degradation products thatresult during hydrolysis (Sofia et al. Journal of Biomedical materialsresearch 2001, 54, 139-148). There also are processing difficulties withpolyesters that can lead to inconsistent hydrolysis rates and tissueresponse profiles. Thus, there is a need for polymeric materials thathave more controllable features such as hydrolysis rates, structure, andmechanical strength, while also being biodegradable and biocompatible.Biological polymeric materials often demonstrate combinations ofproperties which are unable to be reproduced by synthetic polymericmaterials. (Perez-Rigueiro et al. Science, 1998; 70: 2439-2447;Hutmacher D. Biomaterials 2000. 21, 2529-2543). Bone tissue is oneexample; scaffolds for bone tissue regeneration require high mechanicalstrength and porosity along with biodegradability and biocompatibility.

Several studies have shown that BMSCs can differentiate along anosteogenic lineage and form three-dimensional bone-like tissue (Holy etal. J. Biomed. Mater. Res. (2003) 65A:447-453; Karp et al., J.Craniofacial Surgery 14(3): 317-323). However, there are importantlimitations. Some calcium phosphate scaffolds show limited ability todegrade (Ohgushi et al. 1992. In CRC Handbook of bioactive ceramics. T.Yamamuro, L. L. Hench, and J. Wilson, editors. Boca Raton, Fla.: CRCPress. 235-238), or degradation is too rapid (Petite et al. 2000. NatBiotechnol 18:959-963.) Polymeric scaffolds used for bone tissueengineering, such as poly(lactic-co-glycolic acid) or poly-L-lactic acidcan induce inflammation due to acid hydrolysis products, and processingdifficulties can lead to inconsistent hydrolysis rates and tissueresponse profiles (Athanasiou, et al. 1996. Biomaterials 17:93-102;Hollinger et al. 1996. Clin Orthop: 55-65). Difficulties in matchingmechanical properties to support desired function also remain an issue(Harris et al. J Biomed Mater Res 42:396-402).

Studies have also shown that BMSCs can differentiate along chondrogeniclineage and form three-dimensional cartilage-like tissue on biomaterialsubstrates, such as poly(lactic-co-glycolic acid) or poly-L-lactic acid(Caterson et al. 2001. J Biomed Mater Res 57:394-403; Martin et al.2001. J Biomed Mater Res 55:229-235). However, the use of thesescaffolds for cartilage formation present with the same limitations asobserved with their use in bone engineering.

Therefore, in light of the disadvantages in the existing polymers, andeven for more diverse options in degradable polymer systems, thereexists a need for additional biocompatible polymers, particularlypolymers suitable for formation of scaffolds for mechanically robustapplications such as bone or cartilage. The fabrication process for suchscaffolds should be simple, reproducible, and the variables relativelyeasy to control in order to consistently modulate mechanical propertiesand porosity without sacrificing biodegradation.

SUMMARY OF THE INVENTION

The present invention relates to a porous silk fibroin materialcomprising a three-dimensional silk fibroin body having interconnectedpores therein. The pores having a diameter of 50 to 1000 microns. Thepore density is from 20-200 mg/ml, preferably from 40-150 mg/ml.Porosity ranges from 50-99.5%, preferably 70-99%. Most preferably theporosity is above 80%. The density or porosity can be adjusted dependingon the use of material.

Preferably, the material has a compressive modulus of at least 100 kPa.More preferably, the material has a compressive modulus of at least 150kPa. Even more preferably, the material has a compressive modulus of 200kPa. Most preferably, the material has a compressive modulus of 250 kPa.

The porous silk fibroin material of the invention can be used, forexample, as a scaffold for a engineered tissue or as a drug deliverydevice. In tissue engineering, the porosity of the scaffold can beadjusted as to mimic the gradient of cellular densities found in naturaltissue. As such, an organized three-dimensional tissue with a shape andcellular organization substantially similar to that of the tissue invivo is produced.

In one preferred embodiment, a porous silk fibroin material is producedby a process comprising the steps of: (a) forming a silk fibroinsolution comprising silk fibroin in an aqueous salt solution; (b)removing the salt and water from the fibroin solution to form a silkfibroin substance; (c) forming a polymer solution comprising about 5 to35% by weight of the silk fibroin substance in a solvent; (d) contactingthe polymer solution with water-soluble non-toxic particles that areinsoluble in organic solvents and have a diameter between about 50 andabout 1000 microns; (e) placing the polymer solution into a form; (f)removing the solvent from the polymer; (g) contacting said polymersolution with an effective amount of an agent to induce β-sheetstructure and insolubility in aqueous solution; (h) leaching saidpolymer with a solvent in which said particles are soluble and polymeris insoluble to remove said particles from said polymer; and (i) dryingsaid polymer.

Preferably, the solvent in step (c) is selected from the groupconsisting of hexa-fluoro-iso-propanol (HFIP), N-methyl morpholineN-oxide and calcium nitrate-methanol. Most preferably the solvent ishexa-fluoro-iso-propanol (HFIP).

Preferably the solvent is removed from the polymer by sublimination orevaporation.

In another embodiment, the agent to induce β-sheet structure is analcohol and is selected from the group consisting of methanol,2-propanol, or 1-butanol, isoamyl alcohol. The agent may also be asolvent such as chloroform or acetone.

In yet another embodiment, the aqueous salt solution comprises lithiumbromide, lithium thiocyanate, calcium nitrate or related chemicals thatsolubilize the silk.

In one embodiment, the particles are selected from the group consistingof alkali metal and alkaline earth metal halides, phosphates andsulfates, sugar crystals, water-soluble polymer microspheres,polysaccharides and protein microspheres.

In a preferred embodiment, the particles are sodium chloride crystals.

In one embodiment, the polymer used in the method is leached with water.

The invention also provides a porous silk fibroin scaffold materialprepared by the process for producing a porous silk as outlined above.The porous silk fibroin material may be combined with otherbiodegradable polymers.

In one embodiment, the pore density or porosity may be different withinseparate portions of the material. For example, by altering the size ofthe water-soluble particles placed in different positions in the mold,the pore density of the material can be controlled.

The present invention further provides for methods for the production oftissue both in vivo and ex vivo using the porous silk fibroin scaffoldsdescribed herein. Tissues include for example, cartilaginous and bonetissue.

In one embodiment, tissue is produced ex vivo by culturing mammaliancells on a porous silk fibroin scaffold under conditions that areappropriate for inducing the tissue formation. A bioreactor ispreferably used.

In another embodiment, a method for producing tissue in vivo isprovided. The method comprises seeding mammalian cells on a porous silkfibroin scaffold and implanting said scaffold into a patient.

In one aspect of the invention, the mammalian cells that are used forproducing tissue are multipotent cells selected from the groupconsisting of bone marrow stromal cells and adult or embryonic stemcells. In one preferred embodiment, for production of bone tissue, thecells are bone marrow stromal cells.

In one preferred embodiment, the porous silk fibroin scaffold has a3-dimensional structure of a predetermined shape such that the tissueproduced takes the form of the 3-dimensional structure.

In one embodiment, the porous silk fibroin scaffold used in tissueengineering comprises silk fibroin selected from the group consisting ofsilks from silkworms, silks from spiders, silks from geneticallyengineered cells, silks from transgenic plants and animals, silks fromcultured cells, native silk, silk from cloned full or partial sequencesof native silk genes, and silk from synthetic genes encoding silk orsilk-like sequences. Preferably fibroin obtained from Bombyx morisilkworms is used.

In one embodiment, the porous silk fibroin scaffold further comprises anadditive. As used herein, an “additive” is any biologically orpharmaceutically active compound. Additives include, but are not limitedto, peptides, antibodies, DNA, RNA, modified RNA/protein composites,glycogens or other sugars, and alcohols. In one preferred aspect theadditive is a peptide that contains an integrin binding sequence, e.g.RGD. In another preferred aspect the additive is a growth factor.

In another embodiment, the porous silk fibroin scaffold furthercomprises one or more biodegradable polymers selected from the groupconsisting of collagens, polylactic acid or its copolymers, polyglycolicacid or its copolymers, polyanhydrides, elastin, glycosaminoclycans andpolysaccharides.

In still another embodiment, the porous silk fibroin scaffold furthercomprises one or more non-biodegradable polymers selected from the groupconsisting of polyethylene, polystyrene, polymethylmethylcryalte,polyethylene oxide and polyurethanes.

In yet another embodiment, the porous silk fibroin scaffold furthercomprises an agent that enhances proliferation and/or differentiation ofsaid mammalian cells.

In one embodiment, the cells used for tissue engineering are autologouscells; cells derived from the recipient of the bone or cartilageimplant. Autologous cells can be animal cells, e.g. dog, cat, monkey,pig, cow, human, and the like. In one preferred embodiment, theautologous cells are human.

In another embodiment the cells used for tissue engineering are donorcells, derived from a source other than the recipient. In one aspect thedonor cells are allogenic cells, i.e. from the same species as therecipient. In another aspect, the donor cells are derived from adifferent species than the recipient.

The present invention also provides 3-dimensional tissues produced usingthe porous silk fibroin material of the invention.

In yet another embodiment a method for treating a patient in need of animplant is provided. The method comprises implanting a 3-dimensionalporous silk fibroin scaffold of predetermined shape. In one aspect, forexample, the 3-dimensional porous silk fibroin scaffold of predeterminedshape is implanted as a bone substitute or as a cartilage substitute.The 3-dimensional porous silk fibroin scaffold can be seeded withautologous or donor cells prior to implantation. Alternatively, the3-dimensional porous silk fibroin scaffold can be implanted withoutseeding cells.

Unless otherwise defined, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art. Although methods and materials similar or equivalent to thosedescribed herein can be used in the practice or testing of theinvention, the preferred methods and materials are described below. Allpublications, patent applications, patents and other referencesmentioned herein are incorporated by reference. In addition, thematerials, methods and examples are illustrative only and not intendedto be limiting. In case of conflict, the present specification,including definitions, controls.

BRIEF DESCRIPTION OF THE FIGURES

The accompanying figures, which are incorporated in and constitute apart of this specification, illustrate embodiments of the invention and,together with the description, serve to explain the objects, advantages,and principles of the invention.

FIGS. 1A-B shows flow charts for silk processing (FIG. 1A) and silkfabrication (FIG. 1B).

FIGS. 2A-F shows the SEM images of inner and outer structure silkscaffold formed by freeze drying: (2A)-20° C. in 15% methanol (inner)(scale bar; 50 μm), (2B) −20° C. in 15% methanol (outer) (scale bar; 50μm), (2C) −20° C. in 15% 2-propanol (inner) (scale bar; 50 μm), (2D)−20° C. in 15% 2-propanol (outer) (scale bar; 50 μm), (2E) −80° C. in15% methanol (inner) (scale bar; 100 μm) and (2F) −80° C. in 15%2-propanol (inner) (scale bar; 100 μm). (Outer; outside image of thescaffolds, Inner; inside image of the scaffold after fracture in liquidnitrogen).

FIGS. 3A-D shows SEM images of inner and outer structure silk scaffoldby salt leaching and gas foaming methods after methanol treatment (scalebar; 200 μm): (3A) NaCl:silk (10:1 wt %) (inner), (3B) NaCl:silk (10:1wt %) (outer), (3C) NH₄HCO₃:silk (10:1 wt %) (inner) and (3D)NH₄HCO₃:silk (10:1 wt %) (outer).

FIGS. 4A-E shows SEM images of silk scaffold by salt leaching and gasfoaming methods after immersion in alcohols. (4A) salt leached scaffoldimmersed in methanol (scale bar; 200 μm), (4B) salt leached scaffoldimmersed in butanol (scale bar; 100 μm), (4C) gas foamed scaffoldimmersed in methanol (scale bar; 200 μm), (4D) gas foamed scaffoldimmersed in propanol (scale bar; 200 μm), (4E) gas foamed scaffoldimmersed in butanol (scale bar; 100 μm).

FIGS. 5A-D show FTIR of silk fibroin before immersion in alcohol (5A,silk I), after immersion in methanol (5B, silk II), 1-butanol (5C, silkI) and 2-propanol (5D, silk I).

FIGS. 6A and 6B show scanning electron microscopy of collagen (6A) andsilk (6B) scaffolds. Bar length=500 μm.

FIGS. 7A and 7B show biochemical characterization of MSC proliferationon silk, silk-RGD, and collagen scaffolds cultured in control medium for4 weeks in dishes. (7A) DNA content and (7B) MTT activity shown perconstruct. Data represent the average±standard deviation of 3-4scaffolds.

FIGS. 8A-F show scanning electron micrographs of MSCs grown on silk andcollagen scaffolds. Face view of collagen (8A) and silk (8B) matrices,cross section of collagen (8C) and silk (8D) matrices cultured incontrol medium for 4 weeks. Bar=400 μm. (8E, 8F) MSCs cultured for 2weeks on silk scaffolds. Bar=20 μm (8E), 200 μm (8F).

FIGS. 9A and 9B show GAG content expressed as μgchondroitin-sulfate/construct (9A) and chondroitin-sulfate/DNA (g/g)(9B) of MSCs cultured on cross linked-collagen, collagen, silk, andsilk-RGD scaffolds after 2 and 4 weeks in chondrogenic medium. Oneasterisk indicates significant difference to the respective controlgroup (p<0.05). Significantly more GAG was detected on the silk matricescompared to collagen when cultured in chondrogenic medium for 4 weeks(p<0.01). GAG deposition on the silks was similar to cross-linkedcollagen.

FIG. 10 shows collagen type 2 mRNA expression from cells cultured inchondrogenic medium (+) or control medium (−). Data are normalized tothe collagen type 2 expression levels of MSCs at the time of seeding(baseline). Expression was significantly higher for all cells culturedin chondrogenic medium compared to controls (p<0.01). Data represent theaverage±standard deviation of 3-4 constructs.

FIGS. 11A-H show Safranin O staining of sections taken from scaffoldsbased on silk (11A, 11B, 11D) and collagen (11E, 11F, 11H) cultured inthe presence of chondrogenic (11A-11D, 11E-11H) and control media (11D,11H) for 4 weeks. Asterisks indicate polymer, arrows chondrocytes.Immunohistochemistry using an antibody against type 2 collagen is shownin panel C and D. Bar=300 μm (11A, 11D) or 150 μm (11B-11D, 11F-11H).

FIGS. 12A-12H show characterization of BMSCs. (12A) phase-contrastphotomicrographs of passage 2 BMSCs at an original magnification of ×20.(12B-12D) Characterization of chondrogenic differentiation in pelletculture. Pellets were either cultured in chondrogenic medium (12B) orcontrol medium (12C). Pellet diameter is approximately 2 mm, and pelletswere stained with safranin O/Fast Red. (12D) Sulphated GAG/DNA (μg/μg)deposition of passages 1, 3, and 5 BMSCs after 4 weeks. Data representsthe average±standard deviation of 5 pellets. (B) Endoglin expression(CD105) of passage 2 BMSCs. (12F-12H) Characterization of osteoblasticdifferentiation along in pellet culture either treated in osteogenic(12F) or in control medium (12G) and stained according to von Kossa.Pellet diameter is approximately 2 mm. (12H) Calcium deposition/DNA(μg/ng) of passages 1, 3, and 5 BMSCs pellet culture. Passage 1 and 3cells deposited significantly more calcium/DNA than passage 5 cells(p<0.05) and data represents the average±standard deviation of 5pellets.

FIGS. 13A and 13B show biochemical characterization of BMSCdifferentiation on cross linked collagen, collagen, silk, and silk-RGDscaffolds after 2 and 4 weeks in osteogenic culture medium. (13A)Calcium deposition per scaffold and (13B) alkaline phosphatase (AP)activity per scaffold. Data are represented as the average±standarddeviation of 3-4 constructs and asterisks indicate statisticallysignificant differences (p<0.05=*; p<0.01=**).

FIGS. 14A-C show gene expression by cells cultured on collagen, silk,and silk-RGD matrices in osteogenic medium (+) and control medium (−)after 2 and 4 weeks in culture. (14A) expression of bone sialoprotein,(14B) osteopontin, and (14C) BMP-2. Data are shown relative to theexpression of the respective gene in BMSCs prior to seeding and is theaverage±standard deviation of 3-4 constructs.

FIGS. 15A-D show μ-CT images taken from collagen (15A, 15B), andsilk-RGD scaffolds (15C, 15D). Insert in 15C is a magnification from15D. Bar length=1.1 mm.

FIGS. 16A-L show μ-CT images of tissue engineered bone on three silkscaffolds with mean pore sizes of 106 μm (16A, 16D, 16G), 225 μm (16B,16E, 16H), and 425 μm (16C, 16F, 16I). The first row (16A-16C) shows aface view, the second row (16D-16F) a lateral view and the third row(16G-16I) a magnification of A-C. 16J-16L shows scaffolds prior totissue culture of a mean pore size of 106 μm, 225 μm, and 425 μmrespectively.

DETAILED DESCRIPTION OF THE INVENTION

The invention is based upon a surprising finding that silk proteinsprocessed into 3-dimensional architectures provides for suitablebiomaterial and tissue engineering matrices. The present inventionprovides a porous silk fibroin material comprising a 3-dimensional silkfibroin body having pores with a diameter of 50 to 1000 microns. Thepore density can be adjusted depending on the use of material.

The unique properties of silks (mechanical strength, slowbiodegradability, biocompatibility, sequence variants to influenceproperties) provide unique and versatile features from these matrices.The compressive strengths of gas foamed and salt leached porous silkmatrices are comparable with various materials used in tissueengineering to make scaffolds as well as with the strength of corticalbone and cartilage. We discovered that salt leached silk scaffoldsexhibited the best mechanical properties including that they had a highcompressive modulus as well as compressive stress, and they were notbrittle. However, the porous silk fibroin material of the presentinvention is not limited to this method of production.

In a preferred embodiment, scaffold features include high porosity,interconnected pores, biocompatibility, degradability, and pore sizeslarge enough for cell growth (greater than or equal to about 50-100 μm).

The methods useful for silk scaffold fabrication according to thepresent invention include (a) gas foaming, and (b) solventcasting/particulate leaching (Freyman et al. Progress in MaterialsScience 2001; 46:273-282; Mikos et al. Electronic Journal ofBiotechnology, Accurate, 2000; 3: No. 2; Thomson et al. Biomaterials,1998; 19:1935-1943; Widmer et al. Biomaterials, 1998; 19:1945-1955;Zhang R.& Ma P. Journal of Biomedical Material Science, 1999;44:446-455; Nam et al. Journal of Applied Polymer Science, Vol. 81,3008-30021, (2001); Agrawal et al. Journal of Biomedical Materialresources 2001, 55, 141-150; Harris et al. Journal of BiomedicalMaterial Research 1998, 42, 396-402; Hutmacher D. Journal of biomaterialscience polymer science Edn 2001. 12, 107-124). The preferred method formaterial fabrication is leaching.

In the preferred embodiment, the solvent casting/particulate leachingmethod involves mixing a water-soluble porogen, for example, NaCl with aviscous silk polymer solution (Freyman et al. Progress in MaterialsScience 2001; 46:273-282; mikos et al. Electronic Journal ofBiotechnology, Accurate, 2000; 3: No. 2; Agrawal et al. Journal ofBiomedical Material resources 2001, 55, 141-150; Hutmacher D.Biomaterials 2000. 21, 2529-2543; Hutmacher D. Journal of biomaterialscience polymer science Edn 2001. 12, 107-124). The mixture is cast intoa Teflon container where the solvent is evaporated. The result is asalt/polymer composite. The composite is immersed in water to leach outthe salt, resulting in a porous three-dimensional structure.

The silk proteins suitable for use in the present invention ispreferably fibroin or related proteins (i.e., silks from spiders). Asused herein, the term “fibroin” includes selected proteins such as silkfrom spiders. Preferably, fibroin is obtained from a solution containinga dissolved silkworm silk or spider silk. The silkworm silk protein isobtained, for example, from Bombyx mori, and the spider silk is obtainedfrom Nephila clavipes. In the alternative, the silk proteins suitablefor use in the present invention can be obtained from a solutioncontaining a genetically engineered silk, such as from bacteria, yeast,mammalian cells, transgenic animals or transgenic plants.

Silk polymer or silk fibroin solution according to the present inventioncan be prepared by any conventional method known to one skilled in theart. For example, B. mori cocoons are boiled for about 30 minutes in anaqueous solution. Preferably, the aqueous solution is about 0.02MNa₂CO₃. The cocoons are rinsed, for example with water, to extract thesericin proteins and the extracted silk is dissolved in an aqueous saltsolution. Salts useful according to the present invention include,lithium bromide, lithium thiocyanate, calcium nitrate or other chemicalcapable of solubilizing silk. Preferably, the extracted silk isdissolved in about 9-12 M LiBr solution. The salt is consequentlyremoved using, for example, dialysis. See FIG. 1A.

The silk polymer solution is formed by mixing about 5% to 35% by weightof the silk fibroin with a solvent. Preferably about 17% (w/v) of silkis used. Solvents useful according to the present invention includehexa-fluoro-iso-propanyl (HFIP), N-methyl morpholine N-oxide and calciumnitrate-methanol. The preferred solvent is HFIP.

The silk polymer/solvent solution of the present invention is placedinto a form, or mold, containing water-soluble particles, or porogens,that are insoluble in organic solvents. Alternatively, the porogens aremixed with the silk polymer solution prior to placement in the mold. Thediameter of the particles are preferably between about 50-1000 microns.Examples of water-soluble porogens useful according to the presentinvention include, NaCl, alkali metals, alkali earth metal halides,phosphates, and sulfates, sugar crystals, water-soluble microspheres,polysaccharides and protein microspheres.

The solvent is consequently removed using, for example, sublimation orevaporation. The polymer solution is treated with an effective amount ofalcohol to induce β-sheet structure and insolubility in aqueoussolution.

The composite or polymer is immersed in water or other solvent in whichthe particles, or porogens are soluble and polymer is insoluble, toremove the particles, resulting in a porous three-dimensional structure,referred herein to as a “silk fibroin scaffold” or “silk fibroinmaterial.” See FIG. 1B.

In general, regenerated silk fibroin is soluble in water and requiresimmersion in an alcohol or other suitable agent to obtain the β-sheetstructure so that it will be insoluble (Sofia et al. Journal ofBiomedical materials research 2001, 54, 139-148). Therefore, prior tosubmersion into aqueous solutions the silk scaffolds are first soaked ina β-sheet structure inducing agent, such as alcohol to induce the phasetransition to β-sheet structure. The type of a β-sheet structureinducing agent can be used to generate scaffolds with differentproperties. For example, as shown in the following examples, whenmethanol and propanol are used to induce β-sheet structure, theresulting scaffolds are stronger but more brittle and therefore suitablein bone regeneration.

The present invention provides porous silk materials that can be used asscaffolds in tissue regeneration. The porous silk of the presentinvention have strong mechanical properties as well as pore structureand interconnections among the pores, as shown in FIG. 14J-14L. Thedensity and porosity of the silk scaffolds can be measured by liquiddisplacement, similar to published methods (Zhang et al. Journal ofBiomedical Material Science, 1999; 44:446-455). Preferred liquids foruse in liquid displacement measurements include those that do not causeshrinkage or swelling of the silk, such as Hexane.

The density of the silk fibroin material is expressed as:d=W/(V ₂ −V ₃);

where

-   -   d=density    -   W=weight    -   V₁=Volume liquid in a container    -   V₂=total volume of fibroin material and liquid after placement        of material in the container    -   V₃=residual liquid after removal of fibroin material from the        container.

The porosity of the silk fibroin material (ε) is measured by thefollowing formula:ε=(V ₁ −V ₃)/(V ₂ −V ₃).

The scaffold material of the present invention is relatively easy tofabricate and has a controlled porous architecture to allow for cellin-growth and tissue regeneration. The ability to adjust porosityprovides a means by which an organized three-dimensional tissue with ashape and cellular organization substantially similar to that of thetissue in vivo can be produced.

The porous silk scaffolds of the present invention may further includeadditives. The additives can be covalently coupled to the scaffold bymeans known to those skilled in the art. Alternatively, the additive canbe added to media surrounding the scaffold.

Additives suitable for use with the present invention includesbiologically or pharmaceutically active compounds. Examples ofbiologically active compounds include, but are not limited to: cellattachment mediators, such as collagen, elastin, fibronectin,vitronectin, laminin, proteoglycans, or peptides containing knownintegrin binding domains e.g. “RGD” integrin binding sequence, orvariations thereof, that are known to affect cellular attachment(Schaffner P & Dard 2003 Cell Mol Life Sci. January; 60(1):119-32;Hersel U. et al. 2003 Biomaterials. November; 24(24):4385-415);biologically active ligands; and substances that enhance or excludeparticular varieties of cellular or tissue ingrowth. For example, thesteps of cellular repopulation of the scaffold matrix preferably areconducted in the presence of growth factors effective to promoteproliferation of the cultured cells employed to repopulate the matrix.Agents that promote proliferation will be dependent on the cell typeemployed. For example, when fibroblast cells are employed, a growthfactor for use herein may be fibroblast growth factor (FGF), mostpreferably basic fibroblast growth factor (bFGF) (Human RecombinantbFGF, UPSTATE Biotechnology, Inc.). Other examples of additive agentsthat enhance proliferation or differentiation include, but are notlimited to, osteoinductive substances, such as bone morphogenic proteins(BMP); cytokines, growth factors such as epidermal growth factor (EGF),platelet-derived growth factor (PDGF), insulin-like growth factor (IGF-Iand II) TGF-β, and the like. As used herein, the term additive alsoencompasses antibodies, DNA, RNA, modified RNA/protein composites,glycogens or other sugars, and alcohols.

The scaffolds can be shaped into articles for tissue engineering andtissue guided regeneration applications, including reconstructivesurgery. The structure of the scaffold allows generous cellularingrowth, eliminating the need for cellular preseeding. The porouspolymer scaffolds may also be molded to form external scaffolding forthe support of in vitro culturing of cells for the creation of externalsupport organs.

The silk fibroin scaffolds of the present invention may also be mixedwith other biocompatible polymers to form mixed polymer scaffolds. Twoor more biocompatible polymers can be added to the aqueous solutiontogether with the silk polymer. The biocompatible polymer preferred foruse in the present invention is selected from the group comprisingpolyethylene oxide (PEO), polyethylene glycol (PEG), collagen,fibronectin, keratin, polyaspartic acid, polylysine, alginate, chitosan,chitin, hyaluronic acid, pectin, polycaprolactone, polylactic acid,polyglycolic acid, polyhydroxyalkanoates, dextrans, polyanhydrides,polymer, PLA-PGA, polyanhydride, polyorthoester, polycaprolactone,polyfumarate, collagen, chitosan, alginate, hyaluronic acid and otherbiocompatible polymers.

The scaffold functions to mimic the extracellular matrices (ECM) of thebody. The scaffold serves as both a physical support and an adhesivesubstrate for isolated cells during in vitro culture and subsequentimplantation. As the transplanted cell populations grow and the cellsfunction normally, they begin to secrete their own ECM support.

In the reconstruction of structural tissues like cartilage and bone,tissue shape is integral to function, requiring the molding of thescaffold into articles of varying thickness and shape. Any crevices,apertures or refinements desired in the three-dimensional structure canbe created by removing portions of the matrix with scissors, a scalpel,a laser beam or any other cutting instrument. Scaffold applicationsinclude the regeneration of tissues such as nervous, musculoskeletal,cartilaginous, tendenous, hepatic, pancreatic, ocular, integumenary,arteriovenous, urinary or any other tissue forming solid or holloworgans.

The scaffold may also be used in transplantation as a matrix fordissociated cells such to create a three-dimensional tissue or organ.Tissues or organs can be produced by methods of the present inventionfor any species.

A number of different cell types or combinations thereof may be employedin the present invention, depending upon the intended function of thetissue engineered construct being produced. These cell types include,but are not limited to: smooth muscle cells, skeletal muscle cells,cardiac muscle cells, epithelial cells, endothelial cells, urothelialcells, fibroblasts, myoblasts, chondrocytes, chondroblasts, osteoblasts,osteoclasts, keratinocytes, hepatocytes, bile duct cells, pancreaticislet cells, thyroid, parathyroid, adrenal, hypothalamic, pituitary,ovarian, testicular, salivary gland cells, adipocytes, and precursorcells. For example, smooth muscle cells and endothelial cells may beemployed for muscular, tubular constructs, e.g., constructs intended asvascular, esophageal, intestinal, rectal, or ureteral constructs;chondrocytes may be employed in cartilaginous constructs; cardiac musclecells may be employed in heart constructs; hepatocytes and bile ductcells may be employed in liver constructs; epithelial, endothelial,fibroblast, and nerve cells may be employed in constructs intended tofunction as replacements or enhancements for any of the wide variety oftissue types that contain these cells. In general, any cells may beemployed that are found in the natural tissue to which the construct isintended to correspond. In addition, progenitor cells, such as myoblastsor stem cells, may be employed to produce their correspondingdifferentiated cell types. In some instances it may be preferred to useneonatal cells or tumor cells.

Cells can be obtained from donors (allogenic) or from recipients(autologous). Cells can also be of established cell culture lines, oreven cells that have undergone genetic engineering. Pieces of tissue canalso be used, which may provide a number of different cell types in thesame structure.

Appropriate growth conditions for mammalian cells are well known in theart (Freshney, R. I. (2000) Culture of Animal Cells, a Manual of BasicTechnique. Hoboken N.J., John Wiley & Sons; Lanza et al. Principles ofTissue Engineering, Academic Press; 2nd edition May 15, 2000; and Lanza& Atala, Methods of Tissue Engineering Academic Press; 1st editionOctober 2001). Cell culture media generally include essential nutrientsand, optionally, additional elements such as growth factors, salts,minerals, vitamins, etc., that may be selected according to the celltype(s) being cultured. Particular ingredients may be selected toenhance cell growth, differentiation, secretion of specific proteins,etc. In general, standard growth media include Dulbecco's Modified EagleMedium, low glucose (DMEM), with 110 mg/L pyruvate and glutamine,supplemented with 10-20% fetal bovine serum (FBS) or calf serum and 100U/ml penicillin are appropriate as are various other standard media wellknown to those in the art. Growth conditions will vary dependent on thetype of mammalian cells in use and tissue desired.

In one embodiment, methods are provided for producing bone or cartilagetissue in vitro comprising culturing multipotent cells on a porous silkfibroin scaffold under conditions appropriate for inducing bone orcartilage formation. Suitable conditions for the generation of bone andcartilage are well known to those skilled in the art. For example,conditions for the growth of cartilage tissue often comprisenonessential amino acids, ascorbic acid-2-phosphate, dexamethasone,insulin, and TGF-β1. In one preferred embodiment, the nonessential aminoacids are present at a concentration of 0.1 mM, ascorbicacid-2-phosphate is present at a concentration of 50 ug/ml,dexamethasone is present at a concentration of 10 nM, insulin is presentat a concentration of 5 ug/ml and TGF-β1 is present at a concentrationof 5 ng/ml. Suitable conditions for the growth of bone often includeascorbic acid-2-phosphate, dexamethasone, β-glycerolphoasphate andBMP-2. In a preferred embodiment, ascorbic acid-2-phosphate is presentat a concentration of 50 ug/ml, dexamethasone is present at aconcentration of 10 nM, β-glycerolphoasphate is present at aconcentration of 7 mM and BMP-2 is present at a concentration of 1ug/ml.

In general, the length of the growth period will depend on theparticular tissue engineered construct being produced. The growth periodcan be continued until the construct has attained desired properties,e.g., until the construct has reached a particular thickness, size,strength, composition of proteinaceous components, and/or a particularcell density. Methods for assessing these parameters are known to thoseskilled in the art.

Following a first growth period the construct can be seeded with asecond population of cells, which may comprise cells of the same type asused in the first seeding or cells of a different type. The constructcan then be maintained for a second growth period which may be differentin length from the first growth period and may employ different growthconditions. Multiple rounds of cell seeding with intervening growthperiods may be employed.

In one preferred embodiment, tissues and organs are generated forhumans. In other embodiments, tissues and organs are generated foranimals such as, dogs, cats, horses, monkeys, or any other mammal.

The cells are obtained from any suitable donor, either human or animal,or from the subject into which they are to be implanted. As used herein,the term “host” or “subject” includes mammalian species, including, butnot limited to, humans, monkeys, dogs, cows, horses, pigs, sheep, goats,cats, mice, rabbits, rats.

The cells that are used for methods of the present invention should bederived from a source that is compatible with the intended recipient.The cells are dissociated using standard techniques and seeded onto andinto the scaffold. In vitro culturing optionally may be performed priorto implantation. Alternatively, the scaffold is implanted into thesubject, allowed to vascularize, then cells are injected into thescaffold. Methods and reagents for culturing cells in vitro andimplantation of a tissue scaffold are known to those skilled in the art.

Cells can be seeded within the matrix either pre- or post matrixformation, depending on the method of matrix formation. Uniform seedingis preferable. In theory, the number of cells seeded does not limit thefinal tissue produced, however optimal seeding may increase the rate ofgeneration. The number of seeded cells can be optimized using dynamicseeding (Vunjak-Novakovic et al. Biotechnology Progress 1998; Radisic etal. Biotechnoloy and Bioengineering 2003).

It is another aspect of the invention that the 3-dimensional porous silkscaffold, described herein, can itself be implanted in vivo and serve astissue substitute (e.g. to substitute for bone or cartilage). Suchimplants, would require no seeding of cells, but contain an additione.g., RGD, that attracts cells.

In one embodiment, silk matrix scaffolds are seeded with multipotentcells in the presence of media that induces either bone or cartilageformation. Suitable media for the production of cartilage and bone arewell known to those skilled in the art.

As used herein, “multipotent” cells have the ability to differentiateinto more than one cell type in response to distinct differentiationsignals. Examples of multipotent cells include, but are not limited to,bone marrow stromal cells (BMSC) and adult or embryonic stem cells. In apreferred embodiment BMSCs are used. BMSCs are multipotential cells ofthe bone marrow which can proliferate in an undifferentiated state andwith the appropriate extrinsic signals, differentiate into cells ofmesenchymal lineage, such as cartilage, bone, or fat (Friedenstein, A.J. 1976. Int Rev Cytol 47:327-359; Friedenstein et al. 1987. Cell TissueKinet 20:263-272; Caplan, A. I. 1994. Clin Plast Surg 21:429-435; Mackayet al. 1998. Tissue Eng 4:415-428; Herzog et al. Blood. 2003 Nov. 15;102(10):3483-93. Epub 2003 Jul. 31).

The formation of cartilaginous tissue or bone can be monitored by assayswell known to those in the art including, but not limited to, histology,immunohistochemistry, and confocal or scanning electron microscopy (Holyet al., J. Biomed. Mater. Res (2003) 65A:447-453).

Using silk based scaffolds, organized tissue with a predetermined formand structure can be produced either in vitro or in vivo. For example,tissue that is produced ex vivo is functional from the start and can beused as an in vivo implant. Alternatively, the silk based structure canbe seeded with cells capable of forming either bone or cartilage andthen implanted as to promote growth in vivo. Thus, the scaffolds can bedesigned to form tissue with a “customized fit” that is specificallydesigned for implantation in a particular patient. For example,cartilaginous tissue or bone tissue produced by methods of the presentinvention can be used to replace large cartilage or bone defects foundin musculoskeletal disorders and degenerative diseases such asosteoarthritis or rheumatism. Engineered bone and cartilage are alsosuitable for spine and joint replacements such as, elbow, knee, hip orfinger joints or can be used in osteochondral implants.

EXAMPLES Example I Preparation of Scaffolds

Three fabrication techniques; freeze-drying, salt leaching and gasfoaming, were used to form porous three-dimensional silk biomaterialmatrices. Matrices were characterized for morphological and functionalproperties related to processing method and conditions. The porosity ofthe salt leached scaffolds varied between 84-98% with a compressivestrength up to 175±3 KPa and the gas foamed scaffolds had porosities of87-97% and compressive strength up to 280±4 KPa. The freeze-driedscaffolds were prepared at different freezing temperatures (−80 and −20°C.), and subsequently treated with different concentrations (15 and 25%)and hydrophilicity alcohols. The porosity of these scaffolds was up to99% and the maximum compressive strength was 30±2 KPa. Changes in silkfibroin structure during processing to form the 3D matrices wasdetermined by FT-IR and XrD. The salt leached and gas foaming techniquesproduced scaffolds with a useful combination of high compressivestrength, interconnected pores and pore sizes greater than 100 micronsin diameter.

1. Materials

Cocoons of B. mori silkworm silk were kindly supplied by M Tsukada,Institute of Sericulture, Tsukuba, Japan and Marion Goldsmith,University of Rhode Island, USA. Sodium chloride and ammoniumbicarbonate granules were purchased from Sigma. Hexa-fluoro-iso-propanol(HFIP), methanol, 2-propanol, and 1-butanol, were purchased from Aldrichand used without further purification.

Preparation of Regenerated B. mori Silk Fibroin Solutions

B. mori silk fibroin was prepared as follows as a modification of ourearlier procedure (Nam & Young, Science 2001, 81, 3008-30021). Cocoonswere boiled for 30 min in an aqueous solution of 0.02 M Na₂CO₃ thenrinsed thoroughly with water to extract the glue-like sericin proteins.The degummed silk was then dissolved in 9-12 M LiBr solution at roomtemperature yielding a 7-8% (w/v) solution. This solution was dialyzedin water using Slide-a-Lyzer dialysis cassettes (Pierce, MWCO 2000). Thefinal concentration of the aqueous silk fibroin solution was about 2-5.8wt %, which was determined by weighing the remaining solid after drying.The HFIP silk solutions were prepared by dissolving the silk fibroinproduced after lyophilizing the aqueous silk solution into the HFIP.

3D Scaffold Fabrication By Gas Foaming or Salt Leaching Using SilkSolution in HFIP

A viscous silk solution was prepared by dissolving 17% (w/v) silk inHFIP (FIG. 1). Ammonium bicarbonate or sodium chloride particulates(salts particle sizes were 150 to 250 μm) acting as porogens were addedto Teflon disk-shaped molds (diameter and height: 18 and 23 mm). Thesilk/HFIP solution was added into the Teflon molds containing theporogen. The molds were covered to reduce evaporation rate to providesufficient time for more homogenous distribution of the solution in themold. The weight ratios of porogen to silk were adjusted to 10:1 and20:1, respectively. The solvent in the mixture of silk/porogen wasevaporated at room temperature. Just before exposure to water, thecomposite of silk/porogen was immersed in alcohol (methanol, 1-butanolor 2-propanol) for 30 minutes to induce β-sheet structure andinsolubility in aqueous solution. The scaffolds with the ammoniumbicarbonate were immersed in 95° C. water to induce gas foaming forapproximately 10 min until no bubbles were visible and then placed intocold water for an additional 24 hours. The composites that used salt asa porogen were placed into cold water for 24 hours to insure that allthe salt had leached from the matrices to avoid negative impact on cellsin future studies. The water was changed four times a day. The poroussilk scaffolds were air dried and then placed in a vacuum dryer for 24hr.

Scaffold Fabrication By Freeze-Drying Using Aqueous Silk Solutions

A viscous silk precipitate in gel paste form was prepared by adding 15or 25 vol % of methanol or propanol to an aqueous silk solution (5.8 wt%) with gentle mixing (FIG. 1). In this phase of the study 1-butanol wasnot used as a solvent because of its immiscibility in water. A gel pastemixture of the silk/alcohol/water was put into the Teflon disk-shapedmolds. The solution was frozen in dry ice for 2 hour and then placed ina container, which controlled the freezing rate at −1° C./min. Thecontainers were placed into either a −20° C. or −80° C. freezer for 2hrs. The ice/alcohol/silk composite was then lyophilized leaving aporous matrix. In the freeze-drying method, methanol, which is poorsolvent for the silk and induces the formation of β-sheet structure, wasused to induce crystallization and insolubility in water. The alcoholconcentrations were varied between 15 and 25%, selected based on priorstudies.

Scaffold Characterization

SEM

The surface and cross-section morphologies of the scaffolds and poredistributions, sizes, and interconnectivity were observed with a LEOGemini 982 Field Emission Gun SEM. Segments of the outer surface andinner area of the scaffold were prepared by fracture of scaffolds inliquid nitrogen. The specimens were sputter coated with gold. The poresizes were determined by measuring random samples of ten pores from theSEM images using Corel computer software.

Density and Porosity

The density and porosity of the silk scaffolds were measured by liquiddisplacement, similar to the published methods (Zhang et al. Journal ofBiomedical Material Science, 1999; 44:446-455).

Hexane was used as the displacement liquid since it is a non-solvent forsilk and is able to easily permeate through the scaffold and not causeswelling or shrinkage, unlike ethanol. A sample of weight W was immersedin known volume (V1) of hexane in a graduated cylinder. The sample wasleft in the hexane covered for approximately 5 min. In this time thecontents in the cylinder underwent an evacuation-repressurization cycleto force the hexane through the pores. The total volume of the hexaneand the hexane-impregnated scaffold was V2. The volume difference(V2−V1) is the volume of the polymer scaffold. The hexane-impregnatedscaffold was then removed from the cylinder and the residual hexanevolume was recorded as V3. The quantity (V1−V3)—volume of hexane withinthe scaffold—was determined as the void volume of the scaffold. Thetotal volume of the scaffold: V=(V₂−V₁)+(V₁−V₃). The density of the foamis expressed as:d=W/(V ₂ −V ₃)

and the porosity of the foam (ε) was obtained by:ε=(V ₁ −V ₃)/(V ₂ −V ₃)

Mechanical Properties

The compression modulus of the scaffolds was evaluated at roomtemperature on an Instron 881 equipped with a 100N load cell. Cross-headspeed was set at 2 mm/min. Cylinder-shaped samples measuring 8.7 mm indiameter and ranging between 3.2 and 9.5 mm in height were used,according to a modification based on ASTM method F451-95. Thecompressive stress and strain were graphed and the average compressivestrength as well as the compressive modulus and standard deviationdetermined. The elastic modulus is the ratio of the stress to strainwhen deformation is totally elastic; also a measure of stiffness, andthe compressive strength is the maximum engineering stress that may besustained without (Callister et al., 2000 In Materials ScienceEnginering; An Introduction; John Wiley & Sons, Inc, New York, pp114-127).

FTIR

The infrared spectra of the silk fibroin structures were measured with aFT-IR (Bruker Equinox 55) spectrophotometer. The samples were cast on aZnS cell crystal surface directly from silk fibroin solution. The castfilms were treated with alcohols directly. Each spectrum for samples wasacquired in transmittance mode by accumulation of 256 scans with aresolution of 4 cm⁻¹ and a spectral range of 4000-400 cm⁻¹. Thecrystallization effect was investigated on the foams by alcohol type;methanol and 2-propanol. X-ray diffraction was used to determine crystalstructure in the scaffolds from the freeze-drying method. Wide angleX-ray diffraction (WAXD) experiments were conducted on a Bruker D8Discover X-ray diffractometer with general area detector diffraction(GADDS) multiwire area detector; 40 KV, 20 mA and 0.5 mm collimator. Thedistance between the detector and the sample for WAXD was 81.4 mm(CuKα).

Summary of Results and Discussion

Porous silk matrices were formed using three processing methods:freeze-drying, gas foaming, and salt leaching. FIG. 2 shows SEM imagesof freeze-dried scaffolds processed with 15% methanol or 15% 2-propanol,at −20° C. or −80° C. The scaffolds formed highly interconnected andporous structures. The general size of pores observed were small, withdiameters of 50±20 μm. Some of the scaffolds formed two layers, an uppermore porous flake-like layer and a bottom layer that was more condensedand compact that was less brittle. At high concentrations of alcohol(>25%) the silk 3D scaffolds were brittle. The higher the concentrationof silk fibroin in solution the smaller the pores.

Silk fibroin in solution or frozen at low temperature is generally in arandom coil structure based on FTIR analysis. Some aggregation occursduring this process. The silk porous matrices formed at −80° C. resultedin similar morphologies, regardless of the freezing rate or alcohol typeand concentration used in the processing. Pores were small, ˜10 micronsin diameter, and very fine silk aggregates were observed. There was adifference in morphology when comparing the forms formed with methanolversus 2-propanol at −20° C. With 2-propanol a more leaf or sheet likemorphology was observed; while with the methanol smaller spherical poreswere formed. The methanol formed foams were friable and easily broken,unlike the scaffolds formed from 2-propanol. These results likelyreflect differences in miscibility of the two alcohols with water, since2-propanol is not highly miscible, leading to larger pore features.

The glass transition temperature is important in the formation of poreswithin the silk fibroin matrix. The glass transition zone of silk in anaqueous solution is, −20 to −34° C. (Li, M. et al. Journal of AppliedPolymer Science 2001, 79, 2185-2191; Nam, J et al. Journal of AppliedPolymer Science 2001, 81, 3008-30021). The higher the freezingtemperature above the glass transition the longer it will take for iceto form and grow. Therefore, the ice particles will have a greatereffect on the size of the pores. The longer the freezing time then thelarger the pores. Once the temperature is below the glass transitionzone, the pore size is no longer dependent on the temperature. FIG. 2shows that as the temperature was decreased to −80° C. the porediameters decreased to 15±7 μm (Table 1) and the silk fibroin formedsmall aggregates. As the freezing temperature increased the porediameter increased based on SEM. The size of ice particles is dependenton their growth rate and time (Li, M. et al. Journal of Applied PolymerScience 2001, 79, 2185-2191; Nam, J et al. Journal of Applied PolymerScience 2001, 81, 3008-30021; Ishaug-Riley et al. Biomaterials 1998, 19,1405-1412). Above the glass transition the solution is in a stableregion where there is slow growth of ice crystals resulting in amacroporous structure (Li, M. et al. Journal of Applied Polymer Science2001, 79, 2185-2191; Li, M et al. Journal of Applied Polymer Science2001, 79, 2192-2199; Nam, J et al. Journal of Applied Polymer Science2001, 81, 3008-30021). Below the glass transition zone quick freezingshortens the growing time of the ice crystals making it more difficultto grow large ice particles, therefore resulting in an unstable regionwhere microporous interconnected structures are formed (Nam, J et al.Journal of Applied Polymer Science 2001, 81, 3008-30021; Ishaug-Riley etal. Biomaterials 1998, 19, 1405-1412).

The porogen particles used for the salt leaching and gas foamingtechniques were not sieved to obtain a specific porogen size. Therefore,the resulting pores followed a Gaussian distribution. The salt leachedscaffolds (FIGS. 3 and 4) used NaCl as a porogen, which were leached outin water to form the porous matrix. These scaffolds formed pore sizes202±112 μm. The scaffolds varied in structure, some were more ductile,while others were more brittle. The pores were larger in comparison tothose generated by the freeze-drying method. However, the pore structurewas not as highly interconnected; both open and closed pore structureswere observed. A skin layer formed on the surface of the salt leachedscaffolds, which decreased with an increase in porosity.

Scaffolds formed by gas foaming were processed using ammoniumbicarbonate as the porogen. The gas foamed scaffolds with higherporosity were not as strong and flaked apart. SEM images showed a highlyinterconnected open pore morphology with diameters in the range of155±114 μm. The gas foamed process did not leave a skin layer on thesurface of the scaffold. Larger pores could also be formed if largersalt particles were used, the process is only limited by the size of theporogen selected.

Porosity

Matrices with up to 99% porosity were formed depending on the processingmethod: freeze-drying, gas foaming, and salt leaching (Table 1). Thescaffolds prepared by freeze-drying resulted in porosities ˜99%,regardless of the variables studied. With salt leaching and gas foaming,and varying the porogen, (salt to silk ratio of 10:1 or 20:1, wt/wt) theporosities were 84±1% to 98±4% for the salt leached scaffolds and 87±2%to 97±3% for the gas foamed scaffolds.

TABLE 1 Porosity and density of the scaffolds (Ave ± S.D., N = 3) forporosity and density measures. For pore sizes, N = 200. Methods Sampleε(%)¹ d² Pore size (μm) Gas Foaming NH₄HCO₃/Silk (wt %) 10:1 87.0 ± 2.0100 ± 10 155 ± 114 20:1 97.0 ± 1.0 40 ± 5 Salt Leaching NaCl/Silk (wt %)10:1 84.0 ± 2.0 120 ± 2  202 ± 112 20:1 98.0 ± 1.0  40 ± 13 FreezeDrying Alcohol treatment³ (Frozen at −20° C.) 15% Methanol 98.0 ± .0.1020 ± 2 50 ± 20 25% Methanol 99.0 ± 0.01 30 ± 1 15% Propanol 98.0 ± 0.0130 ± 3 25% Propanol 99.0 ± 0.20 30 ± 3 (Frozen at −80° C.) 15% Methanol99.0 ± 0.30 50 ± 3 15 ± 7  15% Propanol 99.0 ± 0.03 40 ± 1 15% Propanol97.0 ± 0.20 30 ± 3 25% Propanol 99.0 ± 0.02 30 ± 5 ¹Porosity ²Density(mg/ml) ³Weight ratio of water in alcoholStructure

FTIR spectra of silk fibroins immersed in methanol, 2-propanol, and1-butanol were observed (FIG. 5). β-sheet structure was observed forsamples immersed in methanol with peaks at 1627.7 and 1524.8 cm⁻¹. SilkI peaks are at 1658.5 and 1524.8 cm⁻¹. Samples immersed in 2-propanolshowed no shift from silk I to β-sheet structure. The samples immersedin 1-butanol showed a slight shift in the peak at 1700 cm⁻¹, indicatingslight β-sheet conformation. X-ray diffraction spectra of freeze-driedsilk scaffolds were analyzed to determine crystallization effects ofalcohol and temperature on the silks scaffolds. The diffraction peaksindicated β-sheet structure for the freeze-dried scaffolds frozen at−80° C. using 25% methanol (FIG. 6). The scaffolds treated withiso-propanol and frozen at −20° C. with 15% methanol had no β-sheetstructure, corroborating the FTIR interpretations.

Mechanical Properties

The structure of the freeze-dried scaffolds was foam like and veryporous. The compressive strength (Table 2) of 15% and 25% methanoltreated foams formed at −20° C. were 10±2 and 10±3 KPa, respectively,with a compressive modulus of 20±1 and 10±3 KPa, respectively. The foamsprocessed using 15 and 25% 2-propanol had a compressive strength of 10±2and 10±3 KPa, respectively, and a compressive modulus of 40±4 and 50±8KPa, respectively. The structures of the foams treated with propanolwere more compact and less porous macroscopically, they did not flakeapart easily. The foams processed using methanol were flakier and easilybroken.

Methanol is a hydrophilic alcohol, miscible with water, and when itcomes into contact with a silk solution it induces the chains of fibrointo interact with each other (Li, M. et al. Journal of Applied PolymerScience 2001, 79, 2185-2191; Li, M et al. Journal of Applied PolymerScience 2001, 79, 2192-2199; Nam, J et al. Journal of Applied PolymerScience 2001, 81, 3008-30021; Magoshi, J. Kobunshi Ronbunshu 1974, 31,765-770; Magoshi, J. et al. Applied Polymer Symposia 1985, 41, 187-204).The fibroin silk chains are then able to organize into a β-sheetstructure. When silk is in the β-sheet structure it is in a crystallineform making it more brittle. 2-Propanol is less hydrophilic thanmethanol and therefore less miscible with water. When added to the silkfibroin solution there is a phase separation. Since propanol is not asmiscible as methanol, less water is drawn from the silk and less β-sheetstructure forms when compared to the methanol treated foams (Li, M. etal. Journal of Applied Polymer Science 2001, 79, 2185-2191; Li, M et al.Journal of Applied Polymer Science 2001, 79, 2192-2199; Nam, J et al.Journal of Applied Polymer Science 2001, 81, 3008-30021; Magoshi, J.Kobunshi Ronbunshu 1974, 31, 765-770; Magoshi, J. et al. Applied PolymerSymposia 1985, 41, 187-204). The foams containing 2-propanol are notcrystalline and are tougher, therefore with a higher compressive modulusthan those processed with methanol. With an increase in theconcentration of methanol there is an increase in compressive modulus.This decrease occurs because of the increase in crystallinity. With ahigher concentration of methanol, more rapid and extensive chainrearrangements occur leading to a higher content of β-sheets within thefoam (Li, M. et al. Journal of Applied Polymer Science 2001, 79,2185-2191; Li, M et al. Journal of Applied Polymer Science 2001, 79,2192-2199).

With the scaffolds prepared at −80° C. with alcohol, methanol and2-propanol, at concentrations of 15 and 25%, there is an increase in thecompressive strength and compressive modulus (Table 2).

TABLE 2 Compressive stress and modulus of the silk scaffolds. AlcoholCompressive Compressive Method Sample Treatment¹ stress (kPa) Modulus(kPa) Gas Foaming NH₄HCO₃/Silk (wt %) 10:1 Methanol 280 ± 4  900 ± 941-Butanol 230 ± 9  500 ± 37 2-Propanol 250 ± 28 800 ± 44 20:1 Methanol250 ± 21 1000 ± 75  1-Butanol 150 ± 8  300 ± 40 2-Propanol 100 ± 11 200± 30 Salt Leaching NaCl/Silk (wt %) 10:1 Methanol  30 ± 10 100 ± 2 1-Butanol 150 ± 14 400 ± 50 2-Propanol 100 ± 20 400 ± 58 20:1 Methanol175 ± 3  450 ± 94 1-Butanol 250 ± 4  490 ± 94 2-Propanol 200 ± 3  790 ±3  Freeze Drying Freeze Temperature −20° C. none 80 ± 1 170 ± 7  15%Methanol 10 ± 2 20 ± 1 25% Methanol 10 ± 3 10 ± 3 15% 2-Propanol 10 ± 240 ± 4 25% 2-Propanol 10 ± 3 50 ± 8 −80° C. none 20 ± 2 220 ± 7  15%Methanol 20 ± 3  90 ± 21 25% 2-Propanol  5 ± 4  90 ± 40 15% Methanol 30± 2 100 ± 1  25% 2-Propanol 20 ± 1 130 ± 1  ¹100% unless otherwiseindicated, weight percent of the alcohol in water.

This increase may be due to the increased freezing rate induced by thetemperature decrease. At this point the foams are able to form onlysmall pores, even though the foams are highly porous, and thecompressive force may distributed differently within this structure ofthe foams.

The gas foamed scaffolds were also treated with alcohols beforecompressive mechanical test. Scaffolds immersed in methanol and propanolhad higher compressive strengths and compressive modulus; 280±4 KPa and900±94 KPa, respectively, and 250±28 and 800±44 KPa, respectively. Whenmechanically crushed, the scaffolds flaked apart and were not as ductileas those produced by the salt leached method. There was a dramaticdifference in compressive stress between the scaffolds with 87% porosityversus 97% porosity. Scaffolds with a high porosity had a lowercompressive strength than the low porosity foams (Table 2).

The salt leached scaffolds prepared using a 10:1 salt to silk ratio hadcompressive strengths from 30 and 150 KPa depending on alcohol treatment(Table 2). Values were higher for the 20:1 samples. These scaffolds wereimmersed in methanol, 1-butanol, and 2-propanol. The scaffolds with 84%porosity were thin and formed thick skin layers, therefore data were notreliable due to the heterogeneous structure.

The gas foamed scaffolds generated a definite compressive yield pointwhere the scaffold was no longer able to maintain its original features.However, the salt leached scaffolds showed characteristics that weremore ductile and sponge-like in behavior. After maximum compression wasreached densification occurred and the scaffold was crushed yet did notflake apart. The cells of the scaffold no longer maintained their shape.

The salt leached scaffolds varied in ductility. The pore structure ofthe salt leached scaffolds was not uniform perhaps due to theevaporation of the organic solvent, HFIP, and the increase in thepolymer concentration of the remaining solution entrapped within thesalt bed. The outer surface of the salt leached scaffolds formed a denseskin layer; common when highly volatile solvents are evaporated from theinterior region of a matrix (Nam, Y. S. et al. Biomaterials 1999, 20,1783-1790). The evaporation of the organic solvent from the surfaceforced the formation of a thick film on the scaffold allowing very fewpores to form for the rest of the solvent to evaporate. This residualsolvent in the matrix function as a plasticizer and makes the polymermore ductile. It often takes up to three days to remove the residualsolvent from these silk matrices, due to this dense skin layer, based onweight changes over time.

The gas foamed scaffolds had a higher compressive strength andcompressive modulus than the salt leached scaffolds. The reasons forthis difference in mechanical properties may be the uniform distributionof the pores within the gas foamed scaffolds which lead to moreconsistent mechanical properties, and/or the pore sizes themselves. Thestress applied to the matrix is concentrated at pore interfaces and ifthe pore distribution is not uniform then the polymer matrix cantypically deform at a lower compressive strength (Harris, L. D et al.Journal of Biomedical Material Research 1998, 42, 396-402; Ehrenstein,G. W. Mechanical Behavior. In Polymeric Materials; Structure,Properties, and Applications; Hanser/Gardener Publications, Inc.,Cincinnati, 2001; pp 167-182). Since the salt leached scaffold consistsof both open and closed pore structures the distribution of thecompressive forces was uneven, causing the scaffold to collapse at lowerstress (Harris, L. D. et al. J. Journal of Biomedical Material Research1998, 42, 396-402).

The mechanical data show the deformation and compression trends of thescaffolds as load is applied. In tissue engineering scaffolds forbone-related applications, matrices are usually designed to hold a loadthat will allow for 1 to 2% strain (Table 3). The gas foamed scaffoldsand the salt leached scaffolds treated with methanol were analyzed. Ofthe scaffolds prepared and evaluated these are the most likely to beused in bone tissue regeneration, therefore the secant modulus, used toassess stress-strain relationships at a specific load level (Kim, S.;Cho, et al. Fibers and Polymers 2001, 2, 64-70), were determined.Compared to porous biodegradable polymeric scaffolds often considered inbone-related tissue engineering studies, the silk porous scaffolds hadsimilar properties at 1 and 2% strain. The gas foamed scaffolds showedthe highest compressive modulus. These data are also comparable to thestrength of cortical bone and cartilage (Table 4).

TABLE 3 Average compressive strength and average compressive modulus ofgas foamed and salt leached scaffold at 1% and 2% strain undercompressive load. (N = 2) Average compressive Average Average moduluscompressive compressive Average compressive Porogen (KPa) at 1% strength(KPa) modulus (KPa) strength (KPa) at 2% Scaffold ratio strain at 1%strain at 2% strain strain Salt 10:1 40 0.40 400 8 Leached 20:1 1600 161200 24 Gas 10:1 500 5 200 16 Foamed 20:1 5000 50 3000 60

TABLE 4 Mechanical properties of three-dimensional porous scaffoldsusing polymeric material, and cortical bone (Suh, H. Yonsei MedicalJournal 1998, 39, 87-96). Compression Compressive MATERIAL Modulus (KPa)strength (KPa) cortical bone 15-30 (GPa) — PLGA^(a)  159 ± 130 —PLGA^(b) 289 ± 25 — PLLA^(c) 242 ± 32 — PLLA^(d) 65 ± 5 — collagen^(e)~150 ~15 chitosan^(f) ~750 ~45 collagen/chitosan^(g) ~500 ~30 silkfibroin 3,000 50 ^(a)poly (d,L-lactic-co-glycolic acid) and NaClparticles were compression molded to form PLGA/NaCl composite. 95%porous scaffold processed using salt leaching method (Harris, L. D.;Kim, B.-S.; Mooney, D. J. Journal of Biomedical Material Research 1998,42, 396-402). ^(b)poly (D,L-lactic-co-glycolic acid) 95% porous scaffoldprocessed using gas foaming method(Harris, L. D.; Kim, B.-S.; Mooney, D.J. Journal of Biomedical Material Research 1998, 42, 396-402).^(c)poly(L-lactic acid) designed using gas foaming method and NH₄HCO₃particles as porogen. The weight ratio of NH₄HCO₃ to PLLA is 10:1 (Nam,Y. S.; Yoon, J. J.; Park, T. G. J. Biomed. Mater. Res. 2000, 53, 1-7).^(d)poly(L-lactic acid) designed using gas foaming method and NH₄HCO₃particles as porogen. The weight ratio of NH₄HCO₃ to PLLA is 20:1(Nam,Y. S.; Yoon, J. J.; Park, T. G. J. Biomed. Mater. Res. 2000, 53, 1-7)..^(e)porous collagen sponge processed by lyophilization (Kim, S et al.Fibers and Polymers 2001, 2, 64-70). ^(f)porous chitosan spongeprocessed by lyophilization (Kim, S et al. Fibers and Polymers 2001, 2,64-70). ^(g)porous sponge made of crosslinked collagen/chitosan,processed by lyophilization (Kim, S et al. Fibers and Polymers 2001, 2,64-70).

Example II Engineering of 3-Dimensional Cartilage Tissue

Materials

Bovine serum, RPMI 1640 medium, Dulbecco's Modified Eagle Medium, basicfibroblast growth factor (bFGF), transforming growth factor-β1 (TGF-β1),Pen-Strep, Fungizone, non essential amino acids, trypsin were from Gibco(Carlsbad, Calif.). Ascorbic acid phosphate, Histopaque-1077, insulin,and dexamethasone were from Sigma (St. Lois, Mo.). Collagen scaffolds(Ultrafoam) were from Davol (Cranston, R.I.). All other substances wereof analytical or pharmaceutical grade and obtained from Sigma. Silkwormcocoons were kindly supplied by M. Tsukada, Institute of Sericulture,Tsukuba, Japan, and Marion Goldsmith, University of Rhode Island).

Scaffold Preparation and Decoration

Cocoons from Bombyx Mori were boiled for 1 hour in an aqueous solutionof 0.02M Na2CO3, and rinsed with water to extract sericins. Purifiedsilk was solubilized in 9M LiBr solution and dialyzed (Pierce, MWCO 2000g/mol) against PBS for 1 day and again against 0.1M MES, 0.5 M NaCl, pH6 buffer for another day. An aliquot of the silk solution was coupledwith GRGDS peptide to obtain RGD-silk. For coupling COOH groups on thesilk were activated by reaction with1-ethyl-3-(dimethylaminopropyl)carbodiimide hydrochloride(EDC)/N-hydroxysuccinimide (NHS) solution for 15 minutes at roomtemperature (Sofia et al. 2001. J Biomed Mater Res 54:139-148). Toquench the EDC, 70 μl/ml β-mercaptoethanol was added. Then 0.5 g/lpeptide was added and left for 2 hours at room temperature. The reactionwas stopped with 10 mM hydroxylamine. Silk solutions were dialyzedagainst 0.1 M 2-(N-morpholino)-ethanesulfonic acid buffer, pH 4.5-5 for1 day. Silk and Silk-RGD solutions were lyophilized and redissolved inhexafluoro-2-propanol (HFIP) to obtain a 17% (w/v) solution. GranularNaCl was weighed in a Teflon container and silk/HFIP solution was addedat a ratio of 20:1 (NaCl/silk). HFIP was allowed to evaporate for 2 daysand NaCl/silk blocks were immersed in 90% (v/v) methanol for 30 minutesto induce a protein conformational transition to β-sheet (Nazarov et al.2003. In Department of Biomedical Engineering. Medford: TuftsUniversity). Blocks were removed, dried and NaCl was extracted in waterfor 2 days. Disk shaped scaffolds (5 mm diameter, 2 mm thick) wereprepared using a dermal punch (Miltey, Lake Success, N.Y.), andautoclaved.

For cross-linking, collagen scaffolds were incubated in 0.1M MES, 0.5 MNaCl, pH=6 buffer for 30 min and subsequently crosslinked with 1.713 gEDC and 0.415 g NHS in 100 ml 0.1M MES, 0.5 M NaCl, pH=6 buffer per gramof collagen (van Wachem et al. 2001. J Biomed Mater Res 55:368-378). Thereaction was allowed to proceed for 4 hrs under gentle shaking and wasstopped by washing for 2 hrs with 0.1 M Na2HPO4. The films were rinsed 4times for 30 min with water. The whole procedure was carried outaseptically.

Human Bone Marrow Stromal Cell Isolation and Expansion.

Total bone marrow (25 cm³, Clonetics, Santa Rosa, Calif.) was diluted in100 ml of isolation medium (5% FBS in RPMI 1640 medium). Cells wereseparated by density gradient centrifugation. Briefly, 20 ml aliquots ofbone marrow suspension were overlaid onto a poly-sucrose gradient(8=1,077 g/cm³, Histopaque, Sigma, St. Louis, Mo.) and centrifuged at800×g for 30 min at room temperature. The cell layer was carefullyremoved, washed in 10 ml isolation medium, pelleted and contaminatingred blood cells were lysed in 5 ml of Pure-Gene Lysis solution. Cellswere pelleted and suspended in expansion medium (DMEM, 10% FBS, 1 ng/mlbFGF) and seeded in 75 cm² flasks at a density of 5×10⁴ cells/cm². Theadherent cells were allowed to reach approximately 80% confluence (12-17days for the first passage). Cells were trypsinized, replated andpassage 2 (P2) cells (80% confluence after 6-8 days), were used for theexperiments.

Tissue Culture

Passage 2 mesenchymal stem cells (7×10⁵ cells) were suspended in 40 μlDMEM and the suspension was used to seed scaffolds, which were prewettedovernight in DMEM. Seeded constructs were placed in 6-well Petri dishesand incubated at 37° C. for 2 hours to allow cell attachment. Tomaintain moisture 30 μl DMEM was added to the scaffolds every 30minutes. Subsequently, 5 ml of chondrogenic or control medium was addedper well. Control medium was DMEM, supplemented with 10% fetal bovineserum and Pen-Strep. Chondrogenic medium included the control medium wasfurther supplemented with 0.1 mM nonessential amino acids, 50 μg/mlascorbic acid-2-phosphate, 10 nm dexamethasone, 5 μg/ml insulin, 5 ng/mlTGF-β1. Medium was exchanged every 2-3 days and constructs wereharvested after 2 and 4 weeks.

Biochemical Analysis and Histology

Constructs were cultured for 6 hours (0 weeks), 1, 2, and 4 weeks incontrol or chondrogenic medium and processed for biochemical analysisand histology. For DNA analysis, 3-4 constructs per group and time pointwere disintegrated after 6 hours (0 weeks), 1, 2 and 4 weeks in cultureusing steel balls and a Microbeater. DNA content (n=3-4) was measuredusing the PicoGreen assay (Molecular Probes, Eugene, Oreg.), accordingto the protocol of the manufacturer. Samples were measuredfluorometrically at an excitation wavelength of 480 nm and an emissionwavelength of 528 nm. For the MTT assay 3-4 constructs were transferredto 2 ml plastic tubes, and 1.5 ml serum free DMEM was added to each wellsupplemented with MTT (0.5 g/l) and incubated in the dark at 37° C., 5%CO2 for 2 hours. Tubes were centrifuged for 10 minutes at 2000 g and thesupernatant was aspirated. Isopropyl alcohol (1.5 ml) was added andconstructs were disintegrated using steel balls and a Microbeater. Tubeswere centrifuged at 2000 g for 10 minutes and absorption was measured at570 nm. Sulphated glycosaminoglycan (GAG) deposition (n=5), was assessedas previously described (Martin et al. 1999. Ann Biomed Eng 27:656-662).Briefly, constructs were frozen, lyophilized for 3 days, weighed, anddigested for 16 hours at 60° C. with proteinase-K. GAG content wasdetermined colorimetrically by dimethylmethylene blue dye binding andmeasured spectrophotometrically at 525 nm (Farndale et al. 1986. BiochimBiophys Acta 883:173-177).

RNA Isolation, Real-Time-Reverse Transcription Polymerase Chain Reaction(Real Time RT-PCR)

Fresh constructs (n=3-4 per group) were transferred into 2 ml plastictubes and 1.5 ml Trizol was added. Constructs were disintegrated usingsteel balls and a Microbeater. Tubes were centrifuged at 12000 g for 10minutes and the supernatant was transferred to a new tube. Chloroform(200 μl) was added to the solution and incubated for 5 minutes at roomtemperature. Tubes were again centrifuged at 12000 g for 15 minutes andthe upper aqueous phase was transferred to a new tube. One volume of 70%ethanol (v/v) was added and applied to an RNeasy mini spin column(Quiagen, Hilden, Germany). The RNA was washed and eluted according tothe manufacturer's protocol. The RNA samples were reverse transcribed incDNA using oligo (dT)-selection according to the manufacturer's protocol(Superscript Preamplification System, Life Technologies, Gaithersburg,Md.). Collagen type 2, MMP1, and MMP2 gene expression was quantifiedusing the ABI Prism 7000 Real Time PCR system (Applied Biosystems,Foster City, Calif.). PCR reaction conditions were 2 min at 50° C., 10min at 95° C., and then 50 cycles at 95° C. for 15 s, and 1 min at 60°C. The expression data were normalized to the expression of thehousekeeping gene, glyceraldehyde-3-phosphate-dehydrogenase (GAPDH).Probes were labeled at the 5′ end with fluorescent dye FAM (VIC forGAPDH) and with the quencher dye TAMRA at the 3′ end. Primer sequencesfor the human collagen type 2 gene were: Forward primer 5′-GGC AAT AGCAGG TTC ACG TAC A-3′ (SEQ ID NO:1), reverse primer 5′-CGA TAA CAG TCTTGC CCC ACT T-3′ (SEQ ID NO:2), probe 5′-CCG GTA TGT TTC GTG CAG CCA TCCT-3′ (SEQ ID NO:3). Primer sequences for the human GAPDH gene were:Forward primer 5′-ATG GGG AAG GTG AAG GTC G-3′ (SEQ ID NO:4), reverseprimer 5′-TAA AAG CCC TGG TGA CC-3′ (SEQ ID NO:5), probe 5′-CGC CCA ATACGA CCA AAT CCG TTG AC-3′ (SEQ ID NO:6). Primers and probes for MMP1 andMMP2 were purchased from Applied Biosciences (Assay on Demand #Hs00233958 ml (MMP1), Hs 00234422 ml (MMP2))

Histology, Immunohistochemistry and Scanning Electron Microscopy

For histology, constructs were dehydrated, embedded in paraffin and cutin 5 μm sections were. To stain for cartilage tissue, sections weretreated with eosin for 1 min, fast green for 5 min, and 0.2% aqueoussafranin O solution for 5 min, rinsed with distilled water, dehydratedthrough xylene, mounted, and placed under a coverslip (Rosenberg, L.1971. J Bone Joint Surg Am 53:69-82).

For immunohistochemistry using a monoclonal antibody against type 2collagen (2B1.5, dilution 1:100, Neomarkers, Fremont, Calif.), paraffinembedded sections of tissue exposed to films were deparrafinized througha series of graded alcohols, treated with protease 2 for 16 min. Theprimary antibody was added to each slide and incubated for 32 minutes atroom temperature in a humidified chamber. The secondary antibody wasapplied, carrying horse-raddish peroxidase and developed according tothe manufacturer's protocol (BenchMark IHC staining module, Ventana,Tucson, Ariz.). Sections were counterstained using hematoxylin for 2minutes.

For SEM, constructs were cut and exposed to Karnovsky's fixative (2%paraformaldehye, 2% glutaraldehyde in 0.1M phosphate Buffer). Afterinitial fixation, constructs were again fixed with 1% osmium tetroxidein 0.1M phosphate buffer for 1 hour. After rinsing with PBS for 15 minthe constructs were dehydrated using a series of graded ethyl alcoholsand dried. Constructs were coated with a thin layer of gold prior toevaluation.

Statistical Analysis

Statistical analysis of data was performed by one-way analysis ofvariance (ANOVA) and Tukey-Kramer procedure for post hoc comparison. pvalues less than 0.05 were considered statistically significant.

Results

Proliferation of Human Bone Marrow Stromal Cells on Silk, Silk-RGD, andCollagen Scaffolds

The proliferation of BMSCs on silk 3D matrices, silk matrices decoratedwith the cell adhesion peptide RGD, and collagen control matrices wasassessed over 4 weeks (FIG. 6). The scaffolds were of porous characterand the lattice was more sheet like for collagen and more sponge likefor the silks (FIG. 6).

A total of 7×10⁵ cells were seeded on each of the matrices, asdetermined by DNA assay and cell counts. About 4-5×10⁵ cells were foundon the matrices 6 hours after seeding (week 0). Cell number droppedapproximately 30% for BMSCs cultured on all substrates between week 0and week 1 (FIG. 7A). After 2 weeks in culture, cell numbers increasedand reached initial levels. Cell proliferation was significantly loweron the collagen matrices compared to silk and silk-RGD constructs(p<0.05). After 4 weeks in culture, cell proliferation continued on thesilk-RGD scaffolds and was significantly higher than on the silkscaffolds (p<0.05). On the silk scaffolds (with and without RGD), totalcell numbers remained stable between weeks 2 and 4 while cell numbersdeclined on the collagen scaffolds where cell numbers were significantlylower compared to the silk or silk-RGD matrices (p<0.001).

The MTT assay was used to determine cell viability (FIG. 7B). During thefirst two weeks, cell viability increased on silk and silk-RGDscaffolds, whereas on collagen the numbers were significantly less thanon the silk (p<0.05) and the silk-RGD versions (p<0.01) at the 2 weektime point. After 4 weeks, cell viability decreased on the silk andsilk-RGD matrices and no statistically significant difference wasobserved for the three biomaterial scaffolds. The loss of cell numberson collagen scaffolds was accompanied by advanced biodegradation of thebiomaterial. Wet weight of matrices seeded with BMSCs and exposed tochondrogenic medium and control medium decreased to 16±6% and 19±6%,respectively, of the initial wet weight. The wet weight of the collagenscaffolds which were not seeded with cells and exposed to control mediumfor 4 weeks decreased to 25±11% of the initial weight. Cross linked (CL)collagen scaffolds remained stable with a wet weight of 90±5% after 4weeks in chondrogenic medium. No change in wet weight for silk andsilk-RGD matrices was observed in chondrogenic (94±9%) or control media(95±13%), or on the silk scaffolds without cells exposed to controlmedium (100±3%) during the 4 weeks of culture.

BMSC proliferation resulted in a formation of cell sheets on thesurfaces of all scaffolds after 4 weeks in culture (FIG. 8A, 8B). Crosssections of collagen (FIG. 8C) and silk (FIG. 8D) scaffolds demonstratedmore homogenous cell growth within the silk matrices compared tocollagens. Polymer degradation in the collagen scaffolds resulted in abreakdown of the support lattice structure, whereas the silks retainedtheir structural integrity throughout the experiment.

BMSCs attached to the silk scaffolds via cellular extensions (FIG. 8E)and cells formed networks after 2 weeks (FIG. 8F) and sheets after 4weeks within the silk scaffold pores (FIG. 8D).

Chondrogenesis

The deposition of sulphated glycosaminoglycans (GAG) on collagen, silk,and silk-RGD matrices after 2 and 4 weeks in culture were assessed (FIG.9). After 2 weeks GAG deposition was the same on all biomaterials. TotalGAG content increased on the silks at 4 weeks (p<0.01 compared to allgroups, and p<0.001 compared to collagen) and decreased on collagenafter 4 weeks compared to 2 weeks (p<0.01). Upon cross-linking, GAGdeposition on the collagen scaffolds was similar and statisticallyindifferent to the silks.

Expression of interstitial collagenase (MMP1) and gelatinase A (MMP2) oncollagen and silk scaffolds cultured in chondrogenic medium wereassessed and normalized relative to the expression of the respectiveconstructs cultured in control medium. Neither MMP 1 nor MMP2 expressionwas significantly increased or decreased in chondrogenic medium ascompared to control medium.

The hallmark of human cartilage differentiation is the expression oftype II collagen. Gene expression of collagen type II by the cellscultured on collagen, silk, and silk-RGD matrices in chondrogenic orcontrol medium was determined after 2 and 4 weeks (FIG. 10). Relative toexpression in control medium, over-expression was observed inchondrogenic medium for all biomaterials after 2 and 4 weeks (p<0.01).Compared to the levels measured at the time of seeding (represented bythe baseline in FIG. 9), increased expression levels were observed formatrices grown in control medium.

In the presence of chondrogenic medium, round or angular shaped cellsresided in cavities or lacunae (FIG. 11). The extracelluar matrix (ECM)was stained with safranin O and the cells were found in the depth of thedeposited matrix. ECM-rich areas within collagen constructs wererelatively small and 100-200 μm in length. In contrast, 500-700 μm longand interconnected ECM-rich zones were observed within the silk andsilk-RGD scaffolds. No differences were observed between the silk andsilk-RGD scaffolds. Cells cultured in control medium had spindle-likeand fibroblastic morphology and did not stain positive with safranin O.The cartilage-like nature of the deposited matrix was corroborated byimmunohistochemistry after 4 weeks. The matrix, which surrounded theround cells stained positive for type 2 collagen on both the collagenand the silks.

Discussion

This work demonstrates that relatively large (500-700 μm) and connectedpieces of human cartilage can be generated in vitro by culturing bonemarrow stromal cells on highly porous and structurally stable silkscaffolds. This is in contrast to the collagen scaffold used in thisstudy, where degradation was observed and the collapsing matrixprevented a similar outcome. At least two scenarios could account tothese observations: (i) the production of matrix metalloproteinases(MMP), leading to enzymatic degradation of the newly deposited cartilagecomponents or (ii) the rapid degradation of the collagen matrix.Diseases such as rheumatoid arthritis, are associated with chondrocytessynthesizing MMPs which collectively degrade all components of cartilage(Cawston et al. 1999. Ann NY Acad Sci 878:120-129). In particular, thecollagenolytic enzyme, e.g. interstitial collagenase also known as MMP1,has been strongly implicated in cartilage destruction (Shlopov et al.1997. Arthritis Rheum 40:2065-2074). Our preliminary results suggestthat the observed GAG depletion on collagen matrices was not associatedwith increased MMP1 or MMP2 transcript levels and no over-expression ofeither enzyme was observed on a RNA level.

The collagen scaffold used in this study was stabilized by physicalmethods, but its durability could be extended for example by usingchemical agents that cross-link the collagen (Carpentier et al. 1969. JThorac Cardiovasc Surg 58:467-483). To understand the significance ofbiodegradation on the decline in GAG content on collagen scaffolds after4 versus 2 weeks, cross linked (CL) collagen scaffolds were prepared, inwhich biodegradation was reduced. GAG content on CL-collagen was similaras compared to the silks after 4 weeks and much higher than on theuntreated collagen (FIG. 9). This suggests that the observed loss in GAGcontent on the natural collagen was mainly due rapid degradation of thescaffold and not to enzymatic effects. Collagen scaffolds prepared byphysical processes resemble more the natural collagen and have lesscytotoxic effects and better biocompatibility when compared tochemically cross-linked materials (van Luyn et al. 1992. J Biomed MaterRes 26:1091-1110). Furthermore, it has been shown that collagenscaffolds prepared by physical processes have less cytotoxic effects andbetter biocompatibility when compared to chemically cross-linkedmaterials (van Luyn et al. 1992. J Biomed Mater Res 26:1091-1110). Forpurposes of cartilage tissue engineering the use of chemicallycross-linked collagen as scaffold material is questionable. In responseto implantation of cross-linked collagens, calcification has beenobserved around the implant (Golomb et al. 1987. Am J Pathol127:122-130). This could potentially interfere with the quality of thetissue engineered cartilage tissue upon implantation. Furthermore silkbased materials can cover a broader range of mechanical properties thancollagen based materials, which may be a substantial advantage to meetphysiological needs at the implantation site (Altman et al. 2003.Biomaterials 24:401-416). Because of these constraints and the focus ofthis study to compare natural polymers, only physically stabilizedscaffolds were further characterized and described in this study.

Human BMSCs were cultivated on highly porous silk scaffolds tofacilitate cell seeding and medium exchange. Approximately 60% of theadministered BMSCs were seeded onto the matrices (FIG. 7A) and cellnumbers declined during the first week. Cell viability, as assessed byMTT activity (FIG. 7B), did not decline within this period, suggestingthat approximately 80% of the seeded cells, determined by DNA, werealive. This conclusion is also reflected by the loss in cell numbersafter 1 week (FIG. 7A). A similar pattern of viability was observed onsilk fibres. In accordance to our findings on 3D scaffolds, BMSCviability on silk fibres decreased within the first week by approx. 30%and after 2 weeks in culture the initial viability was retained (Chen etal. 2003. J Biomed Mater Res in press). This suggests that the observedpattern in viability is more a function of the biomaterial chemistry andnot surface morphology or processing. Seeding strategies can also have asignificant impact on the quality of engineered tissues and variousstatic and dynamic seeding strategies have been described to addressthis issue in three-dimensional scaffolds (Ishaug-Riley et al. 1997. JBiomed Mater Res 36:1-8; Carrier et al. 1999. Biotechnol Bioeng64:580-589; Vunjak-Novakovic et al. 1998. Biotechnol Prog 14:193-202).Generally, dynamic seeding strategies result in more uniform celldistributions in the scaffolds compared to the static seeding strategyused in the present study. The droplet protocol employed here involved a2 hour incubation time in the presence of minimal medium to allow cellattachment, likely contributing to the initial loss of cell viability.Cell distribution upon seeding on the collagen was mainly restricted tothe outer 200-300 μm of the scaffold surface; cells gradually penetratedthe entire scaffold during the incubation period of 4 weeks. Theselimitations are unlikely to be lessened by using dynamic seeding sincethe seeding of the center of the collagen scaffold was prevented byphysical constraints. In contrast, BMSCs were seeded successfullythroughout the silk and silk-RGD scaffolds with almost complete fillingof the scaffold voids as seen in histological sections after 1 week(data not shown). The large diameter of interconnected pores togetherwith the stability of the scaffold geometry due to mechanical propertiesand slow rate of degradation facilitated seeding of the entireconstructs. This was also reflected by the homogenous cell distributionon silk scaffolds in contrast to collagen where cells resided in smallareas even after 4 weeks in culture (FIG. 8C, 8D). Similarly, largeareas of cartilage-like tissue were observed by safranin O staining onsilk and silk-RGD scaffolds, whereas this was confined to relativelysmall areas in the collagen scaffolds (FIG. 11). The cartilage-likenature of the deposited tissue was further confirmed byimmunohistochemistry, using antibodies against type 2 collagen andresulting in a strong binding of the antibody (FIG. 11).

After 4 weeks significantly more cells were present on silk and silk-RGDscaffolds than on collagen scaffolds determined by DNA content (FIG.7A). However, in terms of cell viability no differences could bedetected between the biomaterials (FIG. 7B). This can be due tohindrance in diffusion, at least in part created by dense cell sheetscovering the surface of the scaffolds after 4 weeks (FIG. 8A, 8B). Deadcells, which are detected by the DNA assay (FIG. 7A) but not by the MTTassay (FIG. 7B), could not leave the scaffold interior due to theocclusion by cell sheets on the surfaces.

Some studies have been performed to study the differentiation of humanBMSCs on three-dimensional scaffolds to engineer cartilage has beenreported (for a review see (Hunziker, E. B. 2002. OsteoarthritisCartilage 10:432-463). Two of these studies used poly-L-lactic acid(PLA) or poly(lactic-co-glycolic acid) matrices (Caterson et al. 2001. JBiomed Mater Res 57:394-403; Martin et al. 2001. J Biomed Mater Res55:229-235). Although these polymers have been widely used in studieswith biomaterials, they are known to elicit inflammatory responsesmainly due to the release of acidic hydrolysis products (Athanasiou etal. 1996. Biomaterials 17:93-102). Recently, we have demonstrated invitro and in vivo that the inflammatory response elicited by silk,silk-RGD and collagen is lower than due to PLA films (Meinel et al.,2003). GAG deposition/mg scaffold was approximately an order ofmagnitude less for all biomaterials used in this study than describedfor PLA scaffolds (Martin et al. 2001. J Biomed Mater Res 55:229-235).The study describing chondrogenesis on PLA scaffolds (Martin et al.2001. J Biomed Mater Res 55:229-235) used bovine instead of human BMSCsand twice as much TGF-β1 than used in this study. However deposition ofGAG/DNA was similar compared to the PLA study suggesting that theinduction of chondrogensis/cell was similar in human and bovine cells.The wet weight for the PLA scaffolds was less than for the silkscaffolds and this could contribute to the higher GAG/mg wet (Martin etal. 2001. J Biomed Mater Res 55:229-235).

Stained areas for GAG were larger on silk and silk-RGD scaffolds likelydue to the mechanical stability of the silks. Decoration of silkscaffolds with RGD adhesion sequences resulted in increased levels ofDNA (FIG. 7A), but did not lead to an increase in GAG deposition (FIG.9), as compared to collagen scaffolds.

This study demonstrates by histological evaluation and biochemical andgene expression analysis that chondrogenesis of bone marrow stromalcells can be induced on both collagen and silk scaffolds. However, thebiomechanical properties of collagen were insufficient for cartilagetissue engineering, leading to distributed and unconnected patcheswithin the 3D matrices due to premature degradation and collapse. Incontrast, connected cartilage-like tissue was formed in and on silk andsilk-RGD matrices. When these results are considered in connection withthe unique mechanical properties of silk, the induction of cartilageformation on silk-based materials offers new opportunities forbioengineering of cartilage-like tissue in vitro and the treatment ofcartilage defects in vivo.

Example III Engineering of 3-Dimensional Bone Tissue

Materials

Bovine serum, RPMI 1640 medium, Dulbecco's Modified Eagle Medium (DMEM),basic fibroblast growth factor (bFGF), transforming growth factor-β1(TGF-β1) (R&D Systems, Minneapolis, Minn.), Pen-Strep, Fungizone, nonessential amino acids, trypsin were from Gibco (Carlsbad, Calif.).Ascorbic acid phosphate, Histopaque-1077, insulin, dexamethasone,β-glycerolphosphate were from Sigma (St. Lois, Mo.). Collagen scaffolds(Ultrafoam) were from Davol (Cranston, R.I.). All other substances wereof analytical or pharmaceutical grade and obtained from Sigma. Silkwormcocoons were kindly supplied by M. Tsukada (Institute of Sericulture,Tsukuba, Japan) and Marion Goldsmith (University of Rhode Island,Cranston, R1).

Scaffold Preparation and Decoration

Cocoons from Bombyx Mori (Linne, 1758) were boiled for 1 hour in anaqueous solution of 0.02M Na2CO3, and rinsed with water to extractsericins. Purified silk was solubilized in 9M LiBr solution and dialyzed(Pierce, Woburn, Mass.; MWCO 3500 g/mol) against water for 1 day andagain against 0.1M MES (Pierce), 0.5 M NaCl, pH 6 buffer for anotherday. An aliquot of the silk solution was coupled withglycine-arginine-alanine-aspartate-serine (GRGDS) peptide to obtainRGD-silk. For coupling COOH groups on the silk were activated byreaction with 1-ethyl-3-(dimethylaminopropyl)carbodiimide hydrochloride(EDC)/N-hydroxysuccinimide (NHS) solution for 15 minutes at roomtemperature (Sofia et al. 2001. J Biomed Mater Res 54:139-148). Toquench the EDC, 70 μl/ml β-mercaptoethanol was added. Then 0.5 g/lpeptide was added and left for 2 hours at room temperature. The reactionwas stopped with 10 mM hydroxylamine. Silk solutions were dialyzedagainst water for 2 days. Silk and Silk-RGD solutions were lyophilizedand redissolved in hexafluoro-2-propanol (HFIP) to obtain a 17% (w/v)solution. Granular NaCl was weighed in a Teflon container and silk/HFIPsolution was added at a ratio of 20:1 (NaCl/silk). HFIP was allowed toevaporate for 2 days and NaCl/silk blocks were immersed in 90% (v/v)methanol for 30 minutes to induce a protein conformational transition toβ-sheets (Nazarov et al. 2003. In Department of Biomedical Engineering.Medford: Tufts University). Blocks were removed, dried and NaCl wasextracted out in water for 2 days. Disk shaped scaffolds (5 mm diameter,2 mm thick) were prepared using a dermal punch (Miltey, Lake Success,N.Y.), and autoclaved.

For cross-linking, collagen scaffolds were incubated in 0.1M MES, 0.5 MNaCl, pH=6 buffer for 30 min and subsequently crosslinked with 1.713 gEDC and 0.415 g NHS in 100 ml 0.1M MES, 0.5 M NaCl, pH=6 buffer per gramof collagen (van Wachem et al. 2001. J Biomed Mater Res 55:368-378). Thereaction was allowed to proceed for 4 hrs under gentle shaking and wasstopped by washing for 2 hrs with 0.1 M Na2HPO4. The films were rinsed 4times for 30 min with water. The whole procedure was carried outaseptically.

Iodiniation of GRYDS peptide

To assess the amount of bound RGD to the scaffolds, GRYDS peptide wasiodinated with non radioactive iodine to quantify the amount of boundpeptide in the silk film surface by X-ray photoelectron spectrometer(XPS). The procedure involved first flusing of Sep-Pak C18 reverse phasecartridge (Waters) with 10 ml of a 80:20 mix of methanol:water and thenflushing with 10 ml of 0.1M PBS 0.5 NaCl pH=6 buffer, as previouslydescribed (Sofia et al. 2001. J Biomed Mater Res 54:139-148). ThreeIODO-BEADS (Pierce) were rinsed once with 1 ml of PBS buffer. Eighty mlof PBS and then 10 μl of 3.75 g/L NaI in PBS were added and theactivation was allowed for 5 minutes. Then, 1 ml of 0.1 g/l GRYDSpeptide in PBS was added and the reaction was allowed for 15 min. Beadswere rinsed with PBS and the peptide solution was injected into the C18column followed by elution with 0, 20, 40, and 60% methanol in watersolutions. Fractions were collected and analyzed at 280 nm. Theiodination procedure was repeated with the same peptide throughlyophilization of the desired fractions and resolubilizing in buffer toachieve the desired extent of iodination (1 atom of iodine per moleculeof GRYDS). Iodinated peptide was coupled to silk matrices as describedabove for GRGDS.

Human Bone Marrow Stromal Cell Isolation and Expansion

Total bone marrow (25 cm³, Clonetics, Santa Rosa, Calif.) was diluted in100 ml of isolation medium (5% FBS in RPMI 1640 medium). Cells wereseparated by density gradient centrifugation. Briefly, 20 ml aliquots ofbone marrow suspension were overlaid onto a poly-sucrose gradient(δ=1,077 g/cm³, Histopaque, Sigma, St. Louis, Mo.) and centrifuged at800×g for 30 min at room temperature. The cell layer was carefullyremoved, washed in 10 ml isolation medium, pelleted and contaminatingred blood cells were lysed in 5 ml of Pure-Gene Lysis solution. Cellswere pelleted and suspended in expansion medium (DMEM, 10% FBS, 1 ng/mlbFGF) and seeded in 75 cm2 flasks at a density of 5×10⁴ cells/cm2. Theadherent cells were allowed to reach approximately 80% confluence (12-17days for the first passage). Cells were trypsinized, replated andpassage 2 (P2) cells (80% confluence after 6-8 days), were used for theexperiments.

Pellet Culture

For pellet culture, 2×10⁵ cells were centrifuged at 300 g for 10 min at4° C. The medium was aspirated and replaced with osteogenic,chondrogenic, or control medium. Control medium was DMEM supplementedwith 10% FBS, Pen-Strep and Fungizone, chondrogenic medium was controlmedium supplemented with 0.1 mM nonessential amino acids, 50 μg/mlascorbic acid-2-phosphate, 10 nm dexamethasone, 5 μg/ml insulin, 5 ng/mlTGF β1 and osteogenic medium was control medium supplemented with 50μg/ml ascorbic acid-2-phosphate, 10 nm dexamethasone, 7 mMβ-glycerolphosphate, and 1 μg/ml BMP-2.

Tissue Culture

For cultivation on scaffolds, P2 BMSC (7×10⁵ cells) were suspended inliquid Matrigel® (10 μl) on ice, and the suspension was seeded ontoprewetted (DMEM, overnight) scaffolds. Seeded constructs, placed indishes were placed in an incubator at 37° C. for 15 minutes to allow gelhardening and osteogenic or control medium was added. Half of the mediumwas replaced every 2-3 days. Control medium was DMEM supplemented with10% FBS, Pen-Strep and Fungizone, and osteogenic medium was controlmedium supplemented with 50 μg/ml ascorbic acid-2-phosphate, 10 nmdexamethasone, 7 mM β-glycerolphosphate, and 1 μg/ml BMP-2.

Biochemical Analysis and Histology

Scaffolds were cultured for 2 and 4 weeks in control or osteogenicmedium and processed for biochemical analysis and histology. For DNAanalysis, 3-4 scaffolds per group and time point were disintegratedusing steel balls and a Microbeater. DNA content (n=3-4) was measuredusing the PicoGreen assay (Molecular Probes, Eugene, Oreg.), accordingto the protocol of the manufacturer. Samples were measuredfluorometrically at an excitation wavelength of 480 nm and an emissionwavelength of 528 nm. Sulphated glycosaminoglycan (GAG) deposition forpellet culture (n=5), was assessed as previously described (Martin etal., 1999 Ann Biomed Eng 27:656-662). Briefly, pellets were frozen,lyophilized for 3 days, weighed, and digested for 16 hours at 60° C.with proteinase-K. GAG content was determined colorimetrically bydimethylmethylene blue dye binding and measured spectrophotometricallyat 525 nm (Farndale et al. Biochim Biophys Acta 883:173-177). For totalcalcium content, samples (n=4) were extracted twice with 0.5 ml 5%trichloroacetic acid. Calcium content was determined by a colorimetricassay using o-cresolphthalein complexone (Sigma, St. Louis, Mo.). Thecalcium complex was measured spectrophotometrically at 575 nm. Alkalinephosphatase activity was measured using a biochemical assay from Sigma(St. Louis, Mo.), based on conversion of p-nitrophenyl phosphate top-nitrophenol which was measured spectrophotometrically at 410 nm.

RNA Isolation, Real-Time-Reverse Transcription Polymerase Chain Reaction(Real Time RT-PCR)

Fresh scaffolds (n=3-4 per group) were transferred into 2 ml plastictubes and 1.5 ml Trizol was added. Scaffolds were disintegrated usingsteel balls and a Microbeater. Tubes were centrifuged at 12000 g for 10minutes and the supernatant was transferred to a new tube. Chloroform(200 μl) was added to the solution and incubated for 5 minutes at roomtemperature. Tubes were again centrifuged at 12000 g for 15 minutes andthe upper aqueous phase was transferred to a new tube. One volume of 70%ethanol (v/v) was added and applied to an RNeasy mini spin column(Quiagen, Hilden, Germany). The RNA was washed and eluted according tothe manufacturer's protocol. The RNA samples were reverse transcribedinto cDNA using oligo (dT)-selection according to the manufacturer'sprotocol (Superscript Preamplification System, Life Technologies,Gaithersburg, Md.). Osteopontin, bone sialoprotein, and bone morphogenicprotein 2 gene expression were quantified using the ABI Prism 7000 RealTime PCR system (Applied Biosystems, Foster City, Calif.). PCR reactionconditions were 2 min at 50° C., 10 min at 95° C., and then 50 cycles at95° C. for 15 s, and 1 min at 60° C. The expression data were normalizedto the expression of the housekeeping gene,glyceraldehyde-3-phosphate-dehydrogenase (GAPDH). The GAPDH probe waslabelled at the 5′ end with fluorescent dye VIC and with the quencherdye TAMRA at the 3′ end. Primer sequences for the human GAPDH gene were:Forward primer 5′-ATG GGG AAG GTG AAG GTC G-3′ (SEQ ID NO: 4), reverseprimer 5′-TAA AAG CCC TGG TGA CC-3′ (SEQ ID NO: 5), probe 5′-CGC CCA ATACGA CCA AAT CCG TTG AC-3′ (SEQ ID NO: 6). Primers and probes forosteopontin, bone sialoprotein (BSP), and bone morphogenic protein 2(BMP-2) were purchased from Applied Biosciences (Assay on Demand #Hs00167093 ml (osteopontin), Hs 00173720 ml (BSP), Hs 00214079 ml(BMP-2)).

Histology and Microcomputrized Tomography (μ-CT)

For histology, scaffolds were dehydrated, embedded in paraffin and cutin 5 μm sections were. To stain for cartilage differentiation in thepellet culture (FIG. 12), sections were treated with eosin for 1 min,fast green for 5 min, and 0.2% aqueous safranin O solution for 5 min,rinsed with distilled water, dehydrated through xylene, mounted, andplaced under a coverslip (Rosenberg, L. 1971. J Bone Joint Surg Am53:69-82).

For the visualization of bone distribution, scaffolds were analyzedusing a μCT20 imaging system (Scanco Medical, Bassersdorf, Switzerland)providing a resolution of 34 μm in the face and 250 μm in the crossdirection of the scaffold. A constrained Gaussian filter was used tosuppress noise. Mineralized tissue was segmented from non-mineralizedtissue using a global thresholding procedure. All samples were analyzedusing the same filter width (0.7), filter support (Lysaght et al. 1998.Tissue Eng 4:231-238), and threshold (195/190) as previously described(Muller et al. 1994 Phys Med Biol 39:145-164).

X-Ray Diffraction Measurement (XRD)

X-ray diffraction patterns of scaffolds before and after bone formationwere obtained by means of Bruker D8 Discover X-ray diffractometer withGADDS multiwire area detector. WAXD (wide angle X-ray dirrfaction)experiments were performed employing CuKa radiation (40 kV and 20 mA)and 0.5 mm collimator. The distance between the detector and the samplewas 47 mm.

Statistical Analysis

Statistical analysis of data was performed by one-way analysis ofvariance (ANOVA) and Tukey-Kramer procedure for post hoc comparisonusing SigmaStat 3.0 for Windows. p values less than 0.05 were consideredstatistically significant.

Results

Characterization of Human Bone Marrow Stem Cells

The BMSCs exhibited a spindle shaped and fibroblast-like morphology(FIG. 12A). To assess the differentiation potential of the expandedcells, both chondrogenic and osteogenic inducement was studied. Thechondrogenic potential of P2 cells was evidenced by evenly red stainedareas indicating GAG deposition in the center of the pellets (FIG. 12B).Pellets incubated with either control or osteogenic medium showed no redstaining (insert in FIG. 12C). These findings were corroborated by theanalysis of sulphated GAG content per unit DNA for chondrogenicdifferentiation of P1, P3, and P5 cells (FIG. 12E). No statisticaldifference was observed among the passages. Osteogenic potential of thecells was demonstrated by the spatially uniform deposition of calcifiedmatrix (FIG. 12F). In contrast, pellets incubated in control medium orchondrogenic medium did not show any black or dark brown staining, thusindicating the absence of mineralization. Calcium content per unit DNAwas analyzed to quantify the histological observations for osteogenicdifferentiation for P1, P3, and P5 cells (FIG. 12H). P1 and P3 cellsdeposited significantly more calcium per DNA than P5 cells (p<0.05).BMSCs underwent chondrogenic differentiation only when cultured inchondrogenic medium and osteogenic differentiation only when cultured inosteogenic medium.

Scaffolds, Biochemical Analysis and Real Time RT-PCR

The porosity of the 3D silk matrices was 98%, compressive stress was170±20 kPa and compressive modulus 450±170 kPa (Nazarov et al. 2003. InDepartment of Biomedical Engineering. Medford: Tufts University).

Calcium deposition was analyzed after 2 and 4 weeks in culture (FIG.13A). Calcium content decreased on collagen, which was accompanied withadvanced biodegradation and resulted in a wet weight of approximately17% of the initial weight after 4 weeks as described before (Meinel etal. 2003. Ann Biomed Eng accepted). Cross linked (CL) collagen scaffoldsremained stable with a wet weight of 90±5% after 4 weeks in culture.Calcium deposition was significantly higher when compared to all othergroups (p<0.01) and approximately 3 times more than compared to silk andsilk-RGD and 6 times more than the non CL collagen. A comparison of thenon cross linked scaffold material showed that significantly morecalcium was deposited on collagen than on silk-RGD matrices after 2weeks in culture (p<0.05). In contrast, BMSCs deposited significantlymore calcium on silk (p<0.05) and silk-RGD (p<0.02) scaffolds as oncollagen after 4 weeks. Significantly more calcium was deposited onsilk-RGD than on silk scaffolds after 4 weeks (p<0.05).

Alkaline phosphatase (AP) activity was variable among the groups asindicated by large standard deviations and statistically not significantdifferences (FIG. 13B). AP activity was only significantly higher onsilk-RGD scaffolds when compared to collagen after 4 weeks in culture(p<0.05).

Gene expression of bone sialoprotein (BSP) was upregulated 2000 fold onthe collagen scaffolds cultured in osteogenic medium for 2 weeks whencompared to the expression by BMSCs prior to seeding (represented by thebaseline). When compared to 4 weeks on collagen scaffolds, significantlymore BSP transcript was found on collagen after 2 weeks than after 4weeks in osteogenic medium (p<0.05) (FIG. 14A). Gene expression inosteogenic medium was regulated similarly on the silk, and silk-RGDscaffolds, however the decline in transcripts after 4 weeks whencompared to 2 weeks was not statistically significant. BSP expressionwas higher on silk than on collagen scaffolds after 4 weeks inosteogenic medium (p<0.05). Scaffolds cultured in control medium werenot statistically different from the BMSCs prior to seeding (representedby the baseline).

Osteopontin was over-expressed on all scaffolds cultured in osteogenicmedium and was similar after 2 and 4 weeks in culture, except for silk,where transcript levels increased, however this observation was notstatistically significant (FIG. 14B). Osteopontin expression by allscaffolds cultured in control medium was statistically insignificantfrom the transcript levels measured in BMSCs prior to seeding of eachtime point (represented by the baseline).

Bone morphogenetic protein 2 (BMP-2) expression was 100-150 foldupregulated after 2 weeks on all scaffolds cultured in osteogenic mediumcompared to the BMSCs prior to seeding (FIG. 14C). Fewer transcriptswere found after 4 weeks on collagen and silk-RGD (p<0.05) but not onsilk. The expression of BMP-2 was insignificant for all scaffolds incontrol medium when compared to the expression by BSMCs prior to seeding(represented by the baseline).

Histogenesis of Bone-Like Tissue

After 2 weeks, mineralized spots were observed on the collagen latticeas seen with von Kossa staining. Degradation of the collagen scaffoldwas advanced and the matrix did not provide a stable lattice. Open areasof the lattice were filled with randomly oriented collagen-like bundlesas seen in H&E staining intermingled with fibroblasts. Restricted tocertain locations, enlarged cuboidal cells with an osteoblast-likemorphology were observed. After 4 weeks, calcification on silk scaffoldswas advanced and found adjacent to the scaffold material. Mineralizationoccurred in locations which coalesced and formed clusters of mineralizedmatrix (FIG. 12C). Fewer cells were present after 4 weeks when comparedto 2 weeks and cells with fibroblast morphology along with cells withosteoblast like morphology were observed.

Mineralization on silk was appositional to the scaffold lattice andoccurred in distinct spots (FIG. 12E). The void areas of the matrix werefilled with randomly oriented collagen-like fibres and some fibroblastsand osteoblast like cells were found, predominantly adjacent to thesilk. After 4 weeks, deposition was advanced, especially at the borderzones of the silk. Changes in the extracellular matrix were observedconfined to restricted areas (approx. 50×50 μm) with parallel orientedcollagen bundles surrounded by areas with randomly oriented collagenbundles. Increased numbers of osteoblast like cells with a cuboidal orcolumnar morphology were observed and some of the cells were in contactvia short processes.

On silk-RGD matrices, mineralization occurred in a mode similar to thesilk matrix, appositional to the scaffold lattice. After 2 weeks inosteogenic medium, the void scaffold area was completely filled withconnective tissue, comprised of randomly oriented collagen-like fibres,fibroblasts and cuboidal osteoblast-like cells which were connected viacellular processes. After 4 weeks, mineralization was especiallyadvanced on the silk-RGD scaffolds compared to silk and collagen. Asidefrom the small areas that do not stain, a calcified matrix covered theentire silk-RGD lattice. The void area between the lattices wascompletely filled with extracellular matrix, consisting of paralleloriented collagen bundles, osteoblast like cells, and few cells withfibroblast like morphology. The cells seemed to be interconnected vialong processes.

Bone deposition was imaged after 4 weeks in osteogenic medium withmicro-computerized tomography (μ-CT) of randomly selected collagen andsilk-RGD scaffolds (FIG. 15). Mineralized and unconnected clusters weredistributed mainly at the outer rim of the collagen constructs (FIG.15B). Biodegradation was visible by the concave shape of the scaffoldand a decrease in diameter and thickness by approximately 30% from theoriginal shape (FIG. 15A). No image data were collected for the silkscaffold as mineralization was below the threshold level chosen for theμ-CT imaging. The silk-RGD scaffold showed advanced calcification. Bonedeposition was present at the top and bottom and not in the center ofthe scaffold. The scaffold size remained unchanged during the experimentindicating no substrate degradation during the time frame of theexperiment (FIG. 15C). The calcified rods formed up to 1.2×0.4×0.2 mm(length×width×thickness) interconnected lattices. These interconnectedlattices formed trabecular like geometries, encircling hexagonal voids(not calcified) areas (FIG. 12D, and insert 12C). To identify the bonelike nature of the deposited bone tissue, we compared X-ray diffractionpatterns of poorly crystalline hydroxyapatite (p.c. HA) as predominantlypresent in bone (Harper et al. 1966. Proc Soc Exp Biol Med 122:137-142)and the engineered tissue. Basically, diffracted X-rays from the tissueengineered bone like tissue on silk, collagen, CL-collagen (data notshown) and silk-RGD were the same as of the p.c. HA. The peak observedat 20° in the tissue engineered bone resulted from the silk-RGDscaffold, as determined by a comparison with the XRD pattern of theplain scaffold.

Discussion

In recent years substantial progress has been achieved in the tissueengineering of human autologous bone grafts (Petite et al. 2000. NatBiotechnol 18:959-963; Athanasiou et al. 2000. Tissue Eng 6:361-381;Kale et al. 2000. Nat Biotechnol 18:954-958; Niklason, L. E. 2000. NatBiotechnol 18:929-930; Schoeters et al. 1992. Cell Prolif 25:587-603).To evaluate the potential of silk as a scaffold for bone tissueengineering mesenchymal stem cells (BMSC) derived from human bone marrowwere cultured in three different 3D porous scaffolds (silk, silk-RGD,and collagen) to study osteogenesis under controlled in vitroconditions. BMSCs cultured on the protein scaffolds for up to four weeksformed mineralized bone matrix, and the amount, density, and structureof bone matrix depended on scaffold degradation and scaffold decorationwith RGD.

Human mesenchymal stem cells are a source for autologous bone tissueengineering. They proliferate and differentiate in vitro, can be easilyisolated from bone marrow aspirates, and have a documented potential forosteogenic and chondrogenic differentiation (Friedenstein, A. J. 1976.Int Rev Cytol 47:327-359; Friedenstein et al. 1987. Cell Tissue Kinet20:263-272; Caplan, A. I. 1994. Clin Plast Surg 21:429-435; Pittenger etal. 1999. Science 284:143-147). The osteogenic pathway has been proposedto be the default lineage of this population of cells (Banfi et al.2002. Tissue Eng 8:901-910). The isolated and expanded BMSCs werepositive for the putative BMSC marker CD105/endoglin (Barry et al. 1999.Biochem Biophys Res Commun 265:134-139), and had a capacity forselective differentiation into either cartilage- or bone-forming cells(FIG. 12E). The expanded cells could be induced to undergo eitherchondrogenic or osteogenic differentiation via medium supplementationwith chondrogenic or osteogenic factors, respectively (FIGS. 12B, 12C &12F, 12G). No notable difference in cell differentiation capacity over 3passages in culture was observed, however calcium deposition of passage5 cells was significantly reduced when compared to passage 1 and 3 cells(FIGS. 12D & 12H). In conclusion, the isolated P2 BMSCs used in thisstudy retained osteogenic and chondrogenic differentiation potentialwhich made them a suitable cell source for bone tissue engineering.

Scaffold chemistry and surface modification had a significant impact onmineralization of the matrices. Total calcium deposition per scaffoldwas advanced on collagen scaffolds, but declined after 4 weeks. Thiseffect was correlated with the biodegradation of the collagen scaffold(Meinel et al. 2003. Ann Biomed Eng accepted). Ideally, a scaffold wouldprovide a suitable mechanical match until gradually replaced by thenewly deposited bone. Bone deposition was already present after 2 weeksin culture (FIG. 13A) however, the biodegradation of collagen was toorapid to allow a substantial and stable replacement with newly formedbone. Histological evaluation demonstrated a progression in bonedeposition of the collagen scaffolds between week 2 and 4, althoughtotal calcium per scaffold decreased due to biodegradation. The collagenused in this study did not retain its structure leading to collapsedcollagen fragments intermingled with connective tissue. Presumably, theeroding frame did not allow the bone clusters to connect and, therefore,randomly distributed mineralized clusters were scattered mainly at therim of the scaffold also leading to transport limitations in the centerof the collapsed structure (FIG. 15). To understand the significance ofbiodegradation on the decline in calcium content on collagen scaffoldsafter 4 versus 2 weeks, cross linked (CL) collagen scaffolds wereprepared, in which biodegradation was reduced (Carpentier et al 1969 J.Thorac Cardiovsc Surg 58:467-483). The CL-collagen did not showsubstantial degradation, whereas the natural polymer degraded and wetweight after 4 weeks in culture was about 5 times less as the initialweight. Similarly, calcium content on the CL-collagen was 6 times higheras compared to the untreated and natural polymer (FIG. 13). Thissuggests that the observed loss in calcium content on the naturalcollagen was mainly due rapid degradation of the scaffold. Collagenscaffolds prepared by physical processes resemble more the naturalcollagen and have less cytotoxic effects and better biocompatibilitywhen compared to chemically cross-linked materials (van Luyn et al.1992. J Biomed Mater Res 26:1091-1110). Furthermore silk based materialscan cover a broader range of mechanical properties than collagen basedmaterials, which may be a substantial advantage to meet physiologicalneeds at the implantation site (Altman et al. 2003. Biomaterials24:401-416). Because of these constraints and the focus of this study tocompare natural polymers, only physically stabilized scaffolds werefurther characterized and described in this study.

The introduction of RGD moieties by covalent decoration of the silksurfaces resulted in significantly increased calcium deposition incomparison to non decorated silk or the collagen (FIG. 13A). This is inaccordance with previous studies using silk films decorated with RGD(Sofia et al. 2001. J Biomed Mater Res 54:139-148). Interestingly, boneformation on silk-RGD scaffolds was organized resulting ininterconnected trabeculae of bone like tissue (FIG. 15). The trabeculaeencircle hexagonal void areas, which were in the range of the pore sizesof the silk scaffolds. Histological evaluation corroborated thisevidence, and new bone like tissue was deposited appositionally to thesilk scaffold lattice. These data suggest that the silk scaffoldgeometry may predetermine the geometry of the engineered bone. Calciumdeposition was mainly on the top and bottom of the scaffold. Similar tomost previous studies, it is likely that diffusional limitationsassociated with mass transport have limited successful efforts toengineer compact and continuous bone structures (Ishaug et al. 1997. JBiomed Mater Res 36:17-28; Martin et al. 2001. J Biomed Mater Res55:229-235). A possible avenue to overcome these limitations arebioreactors. Bioreactors support the supply of oxygen, nutrients,metabolites, and regulatory molecules and facilitate mass transfer tothe center of the scaffolds (Freed et al. 2000. In Principles of TissueEngineering. R. P. Lanza, R. Langer, and J. Vacanti, editors. San Diego:Academic Press. 143-156). Since silk biodegradation is slow thescaffolds provide a robust network likely to withstand medium flowconditions (laminar or turbulent flow used in most bioreactors) withoutloss of mechanical integrity.

To characterize the nature of the mineralized tissue, XRD patterns ofengineered bone like tissue were compared to poorly crystallinehydroxyapatite (p.c. HA). From early diffraction measurements it wasconcluded, that bone mineral is a two phase system, one of which wasp.c. HA and the other an amorphous calcium phosphate which makes up lessthan 10% of the mineralized bone (Betts et al. 1975. Proc Natl Acad SciUSA 72:2088-2090). However, more recent studies could not detectamorphous calcium phosphate even in embryonic bone (Bonar et al. 1984. JUltrastruct Res 86:93-99; Grynpas, M. D., et al. 1984. Calcif Tissue Int36:291-301). The similarity in XRD patterns between p.c. HA and theengineered tissue suggested the bone like nature of the depositedinterconnected trabeculae.

Substantial differences in the organization of the extracellular matrixwere observed on the silks between 2 and 4 weeks in osteogenic medium.After 4 weeks dense connective tissue filled the voids of the silk andsilk-RGD lattice in which cuboidal osteoblast like cells were in contactwith one another via long tapering processes. The intercellular spacewas occupied with organized bundles of collagens. This was accompaniedby strong induction of gene expression for a non-collagenous element ofthe extracellular matrix, bone sialoprotein (BSP) (FIG. 14). BSPconstitutes about 15% of the non-collagenous proteins found in themineralized compartment of young bone and supports cell attachmentthrough both RGD dependent and RGD-independent mechanisms, with a highaffinity for hydroxyapatite (Fisher et al. 1983. J Biol Chem258:12723-12727). The strong up-regulation of both genes in response toBMP-2 has been reported in previous studies (Lecanda et al. 1997. J CellBiochem 67:386-396; Zhao et al. 2003. J Dent Res 82:23-27). Theupregulated expression of BSP reflects the strong induction ofextracellular matrix protein production which is seen in thehistological sections. Osteopontin expression was similarly increasedafter 2 and 4 weeks on all scaffolds. Osteopontin regulates celladhesion, migration, survival, and calcium crystal formation, playing arole in biomineralization and early osteogenesis (Butler, W. T. 1989.Connect Tissue Res 23:123-136). Therefore, the increased expression ofosteopontin corroborates the advanced mineralization progress evidentafter 2 weeks in culture.

In summary, tissue engineered organized and trabecular bone-likemorphologies were obtained by using BMSCs cultured in osteogenic mediumon 3D silk scaffolds. This was accompanied by the production of anorganized extracellular matrix. The decoration of silk scaffolds withRGD sequences resulted in increased calcification and a more structuredextracellular matrix. Collagen scaffolds could not generate similaroutcomes due to the rapid rate of degradation. Large tissue engineeredbone with an organized geometry on natural non cross linked polymers canbe engineered by cultivation of BMSCs seeded on silk-RGD scaffolds inconjunction with bioreactors.

Example IV Scaffold Geometry Can Determine Tissue Geometry

Tissue engineered bone was produced on silk material scaffolds that wereprepared as described herein. Several scaffolds were prepared with meanpore sizes of 106 microns, 225 microns, and 425 microns.

Scaffolds were seeded with 5×10⁶ human mesenchymal stem cells andcultured in osteogenic medium as described in Example III for up to 5weeks in a spinner flask. Use of the spinner flask created a turbulentflow of medium around the scaffolds. Bone growth was monitored byμ-computed tomography, the results of which are shown in FIGS. 16A-16I.FIG. 16A-16I shows tissue engineered bone on three silk scaffolds withmean pore sizes of 106 μm (16A, 16D, 16G), 225 μm (16B, 16E, 16H), and425 μm (16C, 16F, 16I). The first row shows a face view, second row alateral view and third row a magnification taken from 16A-16C. The datademonstrates, that plate like bone can be engineered on the scaffoldswith small pore sizes (16A, 16D, 16G), and trabecular structure andorganization can be pre-determined by scaffold geometry. 16I-16L showscaffolds prior to tissue culture of a mean pore size of 106 μm (16J),225 μm (16K), and 425 μm (16L).

All references cited herein and throughout the specification are hereinincorporated by reference in their entirety.

1. A porous silk fibroin material comprising a three-dimensional silkfibroin body having interconnected pores, wherein the pores have adiameter of 155 to 1000 microns, wherein the material has a compressivemodulus of at least 100 kPa.
 2. The porous silk fibroin material ofclaim 1, wherein the pores have a diameter of 155 to 500 microns.
 3. Theporous silk fibroin material of claim 1, wherein the material has acompressive modulus of at least 150 kPa.
 4. The porous silk fibroinmaterial of claim 1, wherein the material has a compressive modulus ofat least 200 kPa.
 5. The porous silk fibroin material of claim 1,wherein the material has a compressive modulus of at least 250 kPa. 6.The porous silk fibroin material of claim 1, wherein the material has aporosity above 80%.
 7. The porous silk fibroin material of claim 1,further comprising at least one additive.
 8. The porous silk fibroinmaterial of claim 7, wherein the additive is a biologically active orpharmaceutically active compound.
 9. The porous silk fibroin material ofclaim 7, wherein at least one additive is a cell growth factor.
 10. Theporous silk fibroin material of claim 9, wherein the cell growth factoris a cytokine.
 11. The porous silk fibroin material of claim 7, whereinat least one additive is a peptide that contains an integrin bindingsequence.
 12. The porous silk fibroin material of claim 7, wherein atleast one additive is a bone morphogenic protein, a basic fibroblastgrowth factor, an epidermal growth factor, a platelet-derived growthfactor, an insulin-like growth factor or a transforming growth factor.13. The porous silk fibroin material of claim 7, wherein the additive isselected from peptides, antibodies, DNA, RNA, modified RNA/proteincomposites, glycogens or other sugars and alcohols.
 14. The porous silkfibroin material of claim 7, wherein the additive is selected fromcollagen, elastin, fibronectin, vitronectin, laminin and proteoglycans.15. A porous silk fibroin material comprising a three-dimensional silkfibroin body having interconnected pores, wherein the pores have a meanpore size of 155±114 to 202±112 micrometers, and wherein the compressivemodulus is 10±3 to 1000±75 kPa.
 16. The porous silk fibroin material ofclaim 15, wherein the material has a porosity above 80%.
 17. The poroussilk fibroin material of claim 15, further comprising at least oneadditive.
 18. The porous silk fibroin material of claim 17, wherein atleast one additive is a cell growth factor.
 19. The porous silk fibroinmaterial of claim 17, wherein at least one additive is a peptide thatcontains an integrin binding sequence.